Matrix composed of a naturally-occurring protein backbone cross linked by a synthetic polymer and methods of generating and using same

ABSTRACT

The present invention relates to biodegradable scaffolds composed of a naturally-occurring protein backbone cross-linked by a synthetic polymer. Specifically, the present invention provides PEGylated-fibrinogen scaffold and methods of generating and using same for treating disorders requiring tissue regeneration.

RELATED APPLICATIONS

This is a continuation-in-part of PCT Patent Application No.PCT/IL2004/001136 filed Dec. 15, 2004, which claims the benefit of U.S.Provisional Patent Application No. 60/530,917 filed Dec. 22, 2003. Thecontents of the above applications are all incorporated by reference.

FIELD AND BACKGROUND OF THE INVENTION

The present invention relates to a matrix composed of anaturally-occurring protein backbone cross-linked by polyethylene glycol(PEG) and, more particularly, to methods of generating and using same intissue regeneration.

Tissue engineering, i.e., the generation of new living tissues in vitro,is widely used to replace diseased, traumatized or other unhealthytissues. The classic tissue engineering approach utilizes living cellsand a basic scaffold for cell culture (Langer and Vacanti, 1993; Neremand Seliktar, 2001). Thus, the scaffold structure attempts to mimic thenatural structure of the tissue it is replacing and to provide atemporary functional support for the cells (Griffith L G, 2002).

Tissue engineering scaffolds are fabricated from either biologicalmaterials or synthetic polymers. Synthetic polymers such as polyethyleneglycol (PEG), Hydroxyapatite/polycaprolactone (HA/PLC), polyglycolicacid (PGA), Poly-L-lactic acid (PLLA), Polymethyl methacrylate (PMMA),polyhydroxyalkanoate (PHA), poly-4-hydroxybutyrate (P4HB), polypropylenefumarate (PPF), polyethylene glycol-dimethacrylate (PEG-DMA),beta-tricalcium phosphate (beta-TCP) and nonbiodegradablepolytetrafluoroethylene (PTFE) provide precise control over themechanical properties of the material (Drury and Mooney, 2003).

Common scaffold fabrication methods are based on foams of syntheticpolymers. However, cell migration into the depth of synthetic scaffoldsis limited by the lack of oxygen and nutrient supply. To overcome suchlimitations, new approaches utilizing solid freeform fabrications andinternal vascular architecture have been developed (Reviewed in SachlosE and Czemuszka J T, 2003; Eur. Cell Mater. 5: 29-39). Likewise,freeze-drying methods are also employed to create uniquethree-dimensional architectures with distinct porosity and permeability.However, creating pores into these materials is an aggressive procedureinvolving the use of toxic reagents which eliminate the possibility ofpre-casting tissue constructs with living cells. Therefore, many of theprefabricated materials are subject to uneven cell seeding andnon-homogeneous populations of cells within the constructs. Furthermore,the materials are generally degraded unevenly during the tissuecultivation process, creating a highly anisotropic tissue with alteredgrowth kinetics.

Scaffolds made of PEG are highly biocompatible (Merrill and Salzman,1983) and exhibit versatile physical characteristics based on theirweight percent, molecular chain length, and cross-linking density(Temenoff J S et al., 2002). In addition, PEG hydrogels are capable of acontrolled liquid-to-solid transition (gelation) in the presence of cellsuspension (Elbert and Hubbell, 2001). Moreover, the PEG gelation (i.e.,PEGylation) reaction can be carried out under non-toxic conditions inthe presence of a photoinitiator (Elisseeff J et al., 2000; Nguyen andWest, 2002) or by mixing a two-part reactive solution of functionalizedPEG and cross-linking constituents (Lutolf and Hubbell, 2003).

However, while the abovementioned synthetic polymers enable precisecontrol over the scaffold material, they often provide inadequatebiological information for cell culture. As a result, these materialsare unsuitable for long-term tissue culture or in vivo tissueregeneration.

On the other hand, naturally occurring scaffolds such as collagen,fibrin, alginate, hyaluronic acid, gelatin, and bacterial cellulose (BC)provide bio-functional signals and exhibit various cellularinteractions. For example, fibrin, a natural substrate of tissueremodeling (Herrick S., et al., 1999), contains several cell-signalingdomains such as a protease degradation substrate (Werb Z, 1999) andcell-adhesion domains (Herrick S., 1999). However, because suchbiological materials exhibit multiple inherent signals (e.g., regulationof cell adhesion, proliferation, cellular phenotype, matrix productionand enzyme activity), their use as scaffolds in tissue regenerationoften results in abnormal regulation of cellular events (Hubbell, 2003).Furthermore, the natural scaffolds are often much weaker afterreconstitution as compared to the strength of the original biologicalmaterial, and little control can be exercised to improve their physicalproperties.

Therefore, the ideal scaffold for tissue engineering should exhibit thestructural characteristics of synthetic materials with thebiofunctionality of natural materials (Leach J B, et al., 2004; Leachand Schmidt, 2005). To this end, several methods of preparing scaffoldwith natural biofunctionality and physical properties of syntheticpolymers have been proposed. Most of these “hybrid” approaches, however,fall short of producing a biomaterial with broad inherentbiofunctionality and a wide range of physical properties; mainly becausethey employ only a single biofunctional element into the materialdesign. For example, prior studies describe the preparation of scaffoldsconsisting of biodegradable elements grafted into the backbone of asynthetic hydrogel network. Hydrogels were prepared from synthetic PEGwhich was cross-linked with short oligopeptides containing enzymaticsubstrates capable of being proteolytically degraded by cell-secretedenzymes [Lutolf et al (2003); Gobin and West (2002)]. Furthermore, toincrease the biofunctionality of such hydrogels, synthetic adhesionmotifs such as the RGD sequences [Lutolf et al (2003)] or VEGF (Seliktaret al; 2004, Zisch A H, et al, 2003; FASEB J. 17: 2260-2. Epub 2003 Oct.16) were grafted into the PEG backbone. However, the use of suchscaffolds (in which PEG is the major component) was limited by theinsufficient bio-feedback and/or long-term cellular responses which areessential for phenotypic stability.

Further attempts to increase the biofunctionality of the scaffoldsincluded the manufacture of genetically-engineered protein-likeprecursors of 100 amino acids, which contain, among other things,several protease substrates and adhesion sites (Halstenberg et al. 2002;Biomacromolecules, 3: 710-23). However, the increased protein precursorssize and the presence of thiol groups required for the PEGylationreaction complicated the purification and solubilization of theprecursors during the scaffold manufacturing process. In addition,similar to the PEG-based biosynthetic materials, thegenetically-engineered protein precursor scaffolds failed to providesufficient biofunctionality to enable long-term stability.

There is thus a wide recognized need for and it would be highlyadvantageous to have biodegradable scaffolds for pomoting tissueregeneration, which are devoid of the above-limitations.

SUMMARY OF THE INVENTION

According to one aspect of the present invention there is provided acomposition-of-matter comprising a naturally occurring protein or abioactive fragment of the protein and at least two synthetic polymerscovalently connected thereto, each of the at least two syntheticpolymers having a functional group being capable of covalently attachingto a syntetic polymer of the at least two synthetic polymers so as toform a scaffold.

According to another aspect of the present invention there is provided ascaffold formed by cross-linking the composition-of-matter.

According to yet another aspect of the present invention there isprovided a scaffold comprising a plurality of protein moleculescovalently attached therebetween so as to form the scaffold via asynthetic polymer having a first part and a second part covalentlyattached therebetween via a chemical moiety being chemically distinctfrom a repeating unit of the polymer.

According to further features in preferred embodiments of the inventiondescribed below, the scaffold is biodegradable.

According to still further features in the described preferredembodiments the synthetic polymer is selected from the group consistingof polyethylene glycol (PEG), Hydroxyapatite/polycaprolactone (HA/PLC),polyglycolic acid (PGA), Poly-L-lactic acid (PLLA), Polymethylmethacrylate (PMMA), polyhydroxyalkanoate (PHA), poly-4-hydroxybutyrate(P4HB), polypropylene fumarate (PPF), polyethylene glycol-dimethacrylate(PEG-DMA), beta-tricalcium phosphate (beta-TCP) and nonbiodegradablepolytetrafluoroethylene (PTFE).

According to still further features in the described preferredembodiments the naturally occurring protein is selected from the groupconsisting of a cell signaling protein, an extracellular matrix protein,a cell adhesion protein, a growth factor and a protease.

According to still further features in the described preferredembodiments the cell signaling protein is selected from the groupconsisting of p38 mitogen-activated protein kinase, nuclear factorkappaB, Raf kinase inhibitor protein (RKIP), Raf-1, MEK, Protein kinaseC (PKC), phosphoinositide-3-kinase gamma, receptor tyrosine kinases(e.g., insulin receptor), heterotrimeric G-proteins [e.g., Galpha(i),Galpha(s) and Galpha(q)], Caveolin-3, and 14-3-3 proteins.

According to still further features in the described preferredembodiments the extracellular matrix protein is selected from the groupconsisting of fibrinogen, Collagen, fibronectin, vimentin,microtubule-associated protein 1D, Neurite outgrowth factor (NOF),bacterial cellulose (BC), laminin and gelatin.

According to still further features in the described preferredembodiments the cell adhesion protein is selected from the groupconsisting of integrin, intercellular adhesion molecule (ICAM) 1, N-CAM,cadherin, tenascin, gicerin, and nerve injury induced protein 2(ninjurin2).

According to still further features in the described preferredembodiments the growth factor is selected from the group consisting ofepidermal growth factor, transforming growth factor-α, fibroblast growthfactor-acidic, fibroblast growth factor-basic, erythropoietin,thrombopoietin, hepatocyte growth factor, insulin-like growth factor-I,insulin-like growth factor-II, Interferon-β, and platelet-derived growthfactor.

According to still further features in the described preferredembodiments the protease protein is selected from the group consistingof pepsin, low specificity chymotrypsin, high specificity chymotrypsin,trypsin, carboxypeptidases, aminopeptidases, proline-endopeptidase,Staphylococcus aureus V8 protease, Proteinase K (PK), aspartic protease,serine proteases, metalloproteases, ADAMTS17, tryptase-gamma, andmatriptase-4.1.

According to still further features in the described preferredembodiments the PEG is selected from the group consisting ofPEG-acrylate (PEG-Ac) and PEG-vinylsulfone (PEG-VS).

According to still further features in the described preferredembodiments the PEG-Ac is selected from the group consisting of PEG-DA,4-arm star PEG multi-Acrylate and 8-arm star PEG multi-Acrylate.

According to still further features in the described preferredembodiments the PEG-DA is a 4-kDa PEG-DA, 6-kDa PEG-DA, 10-kDa PEG-DAand/or 20-kDa PEG-DA.

According to still further features in the described preferredembodiments the synthetic polymer is PEG and whereas the naturallyoccurring protein is fibrinogen.

31 According to still further features in the described preferredembodiments the PEG is selected from the group consisting of PEG-DA,4-arm star PEG multi-Acrylate and 8-arm star PEG multi-Acrylate.

According to still further features in the described preferredembodiments the PEG-DA is a 4-kDa PEG-DA, 6-kDa PEG-DA, 10-kDa PEG-DAand/or 20-kDa PEG-DA.

According to still further features in the described preferredembodiments the fibrinogen is denatured.

According to still further features in the described preferredembodiments a molar ratio between the PEG-DA to the fibrinogen in theunits is 2-400 to 1, respectively.

According to still further features in the described preferredembodiments a concentration of PEG-DA is 3% w/v.

According to still further features in the described preferredembodiments the scaffold further comprising a growth factor.

According to still further features in the described preferredembodiments the growth factor is NGF.

According to still further features in the described preferredembodiments the fibrinogen is whole fibrinogen or fragmented fibrinogen.

According to still further features in the described preferredembodiments the scaffold does not comprise more than 10% unconjugatedform of the syntheric polymer.

According to still further features in the described preferredembodiments the chemical moiety is selected from the group consisting ofaldehyes, acetale, tosyl, tresyl, dichlorotriazine, epoxide, carboxylic,succinimidyle succinate, succinimidyl ester, p-nitrophenyl carbonate,benzotriazolyl carbonate, 2,3,5-trichlorophenyl carbonate, succinimidylecarbonate, pyridildisulphide, maleimide, vinylsulfone, and iodoacetamide.

According to still another aspect of the present invention there isprovided a hydrogel formed from the scaffold.

According to still further features in the described preferredembodiments the naturally occurring protein is whole fibrinogen andwhereas a concentration of the units in the hydrogel is selected from arange of 0.5-35%.

According to still further features in the described preferredembodiments the fibrinogen is fragmented fibrinogen and whereas aconcentration of the units in the hydrogel is selected from a range of0.5-35%.

According to still further features in the described preferredembodiments modulus of elasticity of the hydrogel is in a range of0.02-0.11 kPa for 10-20% polymer.

According to still further features in the described preferredembodiments modulus of elasticity of the hydrogel is in a range of0.01-0.07 kPa for 10-20% polymer.

According to an additional aspect of the present invention there isprovided a method of generating a scaffold comprising:

-   -   (a) covalently attaching a naturally occurring protein or a        bioactive fragment of the protein to at least two synthetic        polymer through a first functional group, each of the at least        two synthetic polymers having a second functional group to        thereby obtain a polymer-protein precursor molecule; and        subsequently    -   (b) cross-linking a plurality of the precursor molecules to        thereby generate the scaffold.

According to still further features in the described preferredembodiments the first and second functional groups are identical.

According to still further features in the described preferredembodiments the method further comprising removing unconjugated form ofthe syntheric polymer prior to step b.

According to yet an additional aspect of the present invention there isprovided a method of inducing in vivo formation of a tissue, the methodcomprising implanting the scaffold in a subject to thereby induce theformation of the tissue.

According to still an additional aspect of the present invention thereis provided a method of inducing ex-vivo formation of a tissue, themethod comprising:

-   -   (a) providing the scaffold; and    -   (b) seeding the scaffold with cells to thereby induce tissue        formation.

According to a further aspect of the present invention there is provideda method of inducing in vivo formation of a tissue, the methodcomprising administering to a subject in need thereof thecomposition-of-matter, the composition being capable of forming thescaffold within the subject, thereby inducing the formation of thetissue in vivo.

According to a further aspect of the present invention there is provideda method of treating a subject having a disorder characterized by tissuedamage or loss, the method comprising:

-   -   (a) providing the scaffold; and    -   (b) implanting the scaffold in the subject to thereby induce        formation of the tissue and treat the disorder characterized by        tissue damage or loss.

According to yet a further aspect of the present invention there isprovided a method of treating a subject having a disorder characterizedby tissue damage or loss, the method comprising:

-   -   (a) providing the scaffold; and    -   (b) seeding the scaffold with cells, and;    -   (c) implanting the scaffold in the subject to thereby induce        formation of the tissue and treat the disorder characterized by        tissue damage or loss.

According to still a further aspect of the present invention there isprovided a method of treating a subject having a disorder characterizedby tissue damage or loss, the method comprising administering to thesubject the composition of matter, the composition being capable offorming a scaffold within the subject, thereby inducing a formation of atissue and treating the disorder characterized by tissue damage or loss.

According to still a further aspect of the present invention there isprovided a composition-of-matter comprising a functional polyethyleneglycol (PEG) attached to denatured fibrinogen or a bioactive fragment ofthe protein.

According to still a further aspect of the present invention there isprovided a kit for inducing tissue regeneration, the kit comprising apackaging material which comprises the composition-of-matter.

According to still a further aspect of the present invention there isprovided a kit for inducing tissue regeneration, the kit comprising apackaging material which comprises the scaffold.

The present invention successfully addresses the shortcomings of thepresently known configurations by providing biodegradable scaffoldsuitable for tissue regeneration applications.

Unless otherwise defined, all technical and scientific terms used hereinhave the same meaning as commonly understood by one of ordinary skill inthe art to which this invention belongs. Although methods and materialssimilar or equivalent to those described herein can be used in thepractice or testing of the present invention, suitable methods andmaterials are described below. In case of conflict, the patentspecification, including definitions, will control. In addition, thematerials, methods, and examples are illustrative only and not intendedto be limiting.

BRIEF DESCRIPTION OF THE DRAWINGS

The file of this patent contains at least one drawing executed in colorphotograph. Copies of this patent with color photograph(s) will beprovided by the Patent and Trademark Office upon request and payment ofnecessary fee.

The invention is herein described, by way of example only, withreference to the accompanying drawings. With specific reference now tothe drawings in detail, it is stressed that the particulars shown are byway of example and for purposes of illustrative discussion of thepreferred embodiments of the present invention only, and are presentedin the cause of providing what is believed to be the most useful andreadily understood description of the principles and conceptual aspectsof the invention. In this regard, no attempt is made to show structuraldetails of the invention in more detail than is necessary for afundamental understanding of the invention, the description taken withthe drawings making apparent to those skilled in the art how the severalforms of the invention may be embodied in practice.

In the drawings:

FIGS. 1 a-b are schematic illustrations of the PEG-fibrinogen hydrogelassembly. FIG. 1 a-PEGylated fibrinogen fragments contain a naturalprotease cleavage site (yellow) and multiple unpaired thiol groups (red)for covalent conjugation of functionalized PEG by Michael-type additionreaction. FIG. 1 b-PEG-fibrinogen hydrogel assembly is accomplished bycovalent cross-linking of unreacted PEG-acrylates to one another,resulting in a hydrogel network of PEGylated fibrinogen. Thecross-linking reaction can be accomplished by free-radicalpolymerization in the presence of very small quantities ofphotoinitiator and low-power light.

FIGS. 2 a-b are Coomassie®-blue staining of SDS-PAGE gels illustratingPEGylated and un-PEGylated Fibrinogen Fragments before and afterPEGylation with functionalized PEG. FIG. 2 a-Electrophoresis ofPEGylated whole fibrinogen in 10% SDS-PAGE. Fibrinogen was incubatedwith either 4-kDa linear fractionalized PEG di-acrylate (PEG-DA; lanes1-3) or with PEG-OH (4 kDa, lane 4). Note the presence of the α-chains(63.5-kDa), β-chains (56-kDa), and γ-chains (47-kDa) of fibrinogen inthe presence of PEG-OH following 12 hours of incubation (lane 4) and theprogressive retardation of the protein fragment in the acrylamide gelfollowing the incubation with 4-kDa linear functionalized PEG for 1 hour(lane 3), 2 hours (lane 2) or overnight (lane 1). Lane 5—molecularweight marker (BIO-RAD SDS-PAGE molecular weight standards, Broad range:200,000-6,000 Da); FIG. 2 b—Electrophoresis of PEGylated cyanogenbromide (CNBr)—cleaved fibrinogen fragments in 20% SDS-PAGE. Fibrinogenwas treated with CNBr for 12 hours as described under Materials andExperimental Methods of Example 1 of the Examples section which followsand the cleaved fibrinogen fragments were incubated for 1 hour in thepresence (lanes 1 and 2) or absence (lane 3) of linear fractionalizedPEG-DA. Note the highly PEGylated fibrinogen species (i.e., the highmolecular weight fragments) present following 1 hour incubation withlinear PEG-DA 6-kDa (lane 1) and 4-kDa (lane 2). Lane 4—molecular weightmarker.

FIGS. 3 a-b are graphs illustrating stress-strain characteristics ofPEG-based hydrogels. The slope of the stress-strain curve represents themodulus of the material. The dependency of the modulus [expressed inunits of kilo Pascal (kPa)] on the length of PEG used was measured usingan Instron™ 5544 single column material testing system with the Merlinsoftware (www.instron.com). FIG. 3 a—behavior of 20% polymer (w/v)PEG-PEG hydrogels; FIG. 3 b—behavior of PEG-fibrinogen hydrogels. Notethat material properties (modulus) are dependent on the molecular weightof the PEG (FIG. 3 a). Also note that while in the PEG-PEG hydrogels themodulus of the hydrogel is directly proportional to molecular weight ofthe PEG (FIG. 3 a), in the PEG-fibrinogen hydrogels similar modulusvalues were observed using the 4-kDa and the 6-kDa PEG. Aside from thisexception, the stress-strain behavior of the PEG-fibrinogen hydrogelswas identical to that of PEG-only controls.

FIGS. 4 a-c are graphs illustrating the ability to manipulate themechanical properties of the PEG-fibrinogen hydrogels by altering thepercent of polymer used (FIG. 4 a), the PEG molecular weight used (FIG.4 b) and the polymer/protein composition (FIG. 4 c). Introduction ofvariations in these parameters enables the generation of hydrogels witha broad range of properties, including differing elastic moduli(measured in kPa). FIG. 4 a shows the effect of increasing the percentpolymeric composition of 4-kDa PEGylated fibrinogen on the hydrogelelastic modulus; note that the elastic modulus of PEG-fibrinogenhydrogels (cleaved and whole protein) proportionally increases withincreasing percent polymeric composition (also true for 6-kDa and 20-kDaPEGylated fibrinogen hydrogels). FIG. 4 b demonstrates the effect ofincreasing the PEG molecular weight (using 20% polymer) on the hydrogelelastic modulus; note that the molecular weight of the PEG constituenthas a direct proportional effect on the elastic modulus of the hydrogels(also true for 10% and 15% PEG-fibrinogen hydrogels). FIG. 4 c shows theeffect of increasing the ratio of PEGylated fibrinogen to additionalPEG-DA in the composition of the hydrogels (using 15% polymer and wholePEGylated fibrinogen). Note that the relative amount of PEGylatedfibrinogen and free, unbound PEG-DA in the precursor solution directlyimpacts the stiffness of the hydrogel after the free-radicalpolymerization; the graph shows the increase in stiffness of thePEG-fibrinogen hydrogel as a function of the addition of free PEG-DA.Note that hydrogels made only with PEG-DA (no fibrinogen) are alwaysstiffer in comparison to PEGylated fibrinogen hydrogels. All error barsshow standard deviations from the mean.

FIGS. 5 a-b are graphs illustrating protease-mediated degradation ofPEG-fibrinogen hydrogels. The release of colorimetric fibrinogenfragments from the hydrogels was used to assess protease-mediateddegradation following 30 minutes incubation with the proteasecollagenase or trypsin. FIG. 5 a-illustrates protease-mediateddegradation in the presence of 0.5 mg/ml Collagenase or 0.05 mg/mltrypsin solution after 30 minutes incubation at 37° C. and mildagitation. Note that the amount of degradation taking place bytrypsin-mediated proteolysis is significantly affected by the molecularweight of the grafted PEG constituent [i.e., more degradation in thecase of the high molecular weight (20 kDa) PEG], whereas the amount ofdegradation taking place by Collagenase-mediated proteolysis issignificantly affected by the fibrinogen constituent (i.e., moredegradation in the case of cleaved fibrinogen). FIG. 5 b—illustratesprotease-mediated degradation as a function of the proteaseconcentration. PEG-fibrinogen hydrogels consisting of 15% PEG (6 kDamolecular weight) and whole fibrinogen were subjected to Trypsin andCollagenase degradation as described hereinabove. Note thedose-dependent response of degradation products as a function ofprotease concentration. All degradation data is normalized with thecolorimetric values of the fully degraded hydrogels in its respectiveprotease solution.

FIGS. 6 a-e are phase-contrast microscopic images of vascular cellcultures in PEG-based Hydrogels. FIG. 6 a—A monolayer of bovine aorticendothelial cells after growing for 24 hours on a PEG-fibrinogenhydrogel surface (4-kDa PEG, 10% polymer); FIGS. 6 b-c—Bovine aorticsmooth muscle cells after growing three-dimensionally for 24 hoursinside a PEG-fibrinogen hydrogel (4-kDa PEG, 10% polymer). Note thethree-dimensional attachment and spreading of the cells in the twoseparate z-slices of the gel (FIGS. 6 b and c); FIG. 6 d—The monolayerof bovine aortic endothelial cells as in FIG. 6 a after growing for 24hours on a PEG-PEG control hydrogel (4-kDa PEG, 10% polymer); FIG. 6e—The aortic smooth muscle cells as in FIGS. 6 b-c after growing insidea PEG-PEG hydrogel (4 kDa PEG, 10%). Note the adhesion and spreading ofthe endothelial cells (FIG. 6 a), as well as the smooth muscle cells(FIGS. 6 b-c) when grown on the PEG-fibrinogen hydrogels but not inPEG-PEG control hydrogels (FIGS. 6 d-e). In the absence of a proteolyticsubstrate in the PEG-PEG control hydrogels, note that the encapsulatedsmooth muscle cells are devoid of cell extensions (FIG. 6 e). All imageswere digitally acquired at 200× magnification using a phase-contrastmicroscope.

FIGS. 7 a-b are histological cross-sections of similar cell-seededPEG-based hydrogels imaged in FIG. 6. Bovine aortic smooth muscle cellswere cultured for 48 hours in a PEG-fibrinogen hydrogel (4 kDa PEG, 10%polymer, FIG. 7 a) or a PEG-PEG control hydrogel (4 kDa, 10%); thehydrogel specimens were cut into 7-μm thick sections and stained withhematoxylin-eosin (H&E) which depicts the cell nucleus in dark purplestaining. Note the spreading of cells within the gel network in thePEG-fibrinogen hydrogel (FIG. 7 a) as opposed to the encapsulated cellsshown within the PEG-PEG control hydrogel, which are proteolyticallynon-degradable.

FIGS. 8 a-d are photographs illustrating the formation of critical sizerat tibia defect model. FIG. 8 a—illustrates the insertion of externalfixation into the rat tibia; FIG. 8 b—illustrates the exposure of thetibia; FIG. 8 c—illustrates the excision of a 7-mm wide section of thetibia; FIG. 8 d—illustrates the insertion of a 3-mm diameter precastsolid Gelrin™ cylinder (which is made of the PEG-fibrinogen hydrogel ofthe present invention) into the defect site.

FIGS. 9 a-b are x-ray images depicting the formation of new bone tissueinside the critical size rat tibia defect at 5-weeks post-operation andGelrin™ implantation. The formation of a periosteal callus is clearlyseen in the Gelrin™-treated animal (FIG. 9 b, arrow) as compared withthe control, untreated animal (FIG. 9 a, arrow points to missing bone).Although the X-ray image of the control animal shown was taken 3-weekspost-operatively, similar results were obtained at 5 weekspost-operation (data not shown).

FIG. 10 is a photomicrograph of a rat tibia histological section stainedwith hematoxylin-eosin (H&E) showing new bone formation in the gapregion of Gelrin™-implanted rat tibia at 5-weeks following the inductionof a critical size defect. Note the apparently almost-normalcompacta-type bone, compatible with the usual cortex of the rat bone,the Haversian systems with a small central canal containing bloodvessels (small arrow), and the concentrically organized lamellae. Alsonote the normal osteocytes (OST), one cell to a lacunae, such that thereis no indication of a recent remodeling event. Even though a polarizedphotomicrograph is not available, the bone appears to be everywherelamellar-patterned.

FIG. 11 is a photomicrograph of a rat tibia histological section stainedwith hematoxylin-eosin (H&E) showing osteonal healing with acartilaginous island in the gap region of Gelrin™-implanted rat tibia at5-weeks following the induction of a critical size defect. Note thatwhile the cartilage itself is avascular, a vascular invasion is seen inthe upper and lower borders of the cartilage (short arrows), indicativeof endochondral ossification. Furthermore, at the upper border, thefibro-vascular tissue invading the cartilage is accompanied by cuboidalosteoblasts (OST), which indicates that the deposition of primary boneis taking place. Also note the osteoblastic rimming of the bone(evidence of ongoing osteogenesis) which is seen at the lower part,partly by cuboidal osteoblasts (OST) and partly by bone lining cells(BLC). The features in this image are characteristic of normal bonerepair; the like of which one encounters in the gap of a fractured bonewhich has been fixated without close contact.

FIG. 12 is a photomicrograph of a rat tibia histological section stainedwith hematoxylin-eosin (H&E) showing a remnant of Gelrin™ in the gapregion of Gelrin™-implanted rat tibia at 5-weeks following the inductionof a critical size defect. The image shows the degraded Gelrin™ (star)fibrotically encapsulated, and regions of basic multicellular units(BMU) surrounding. Note the numerous newly formed osseous trabeculae(long arrows) and osseous trabeculae separated by wide tracks ofcellular fibrous tissue (double pointed arrows). It is noteworthy tomention that similar features are encountered during gap healing of abone fracture without contact of the bone edges.

FIG. 13 is a high magnification photomicrograph of a rat tibiahistological section stained with hematoxylin-eosin (H&E) showing aremnant of Gelrin™ in the gap region of Gelrin™-implanted rat tibia at5-weeks following the induction of a critical size defect. This imageshows in greater detail the dense connective matrix and formation of newbone near the residual Gelrin™ of a rat tibia 5-weeks followingimplantation. Short arrows indicate the regions of residual Gelrin™surrounded by dense connective matrix and formation of new bone[asterisk (*)]. The cut edge of the bone with BMU (on the right lowercorner) indicates the remodeling activity of mature bone in which theBMU prepares the way for new bone formation. The appearance of severalsmall areas of residual Gelrin™ suggests that this tissue sectionrepresents the center of the gap. Note the presence of fibrous tissuewithout an inflammatory response around the two round collections ofbroken-up Gelrin™ (short arrows), demonstrating the biocompatiblecharacteristics of the Gelrin™ hydrogel.

FIG. 14 is a photomicrograph depicting new bone formation in the gapregion of a Gelrin™-treated animal. Shown is a histological section of arat tibia 5 weeks following the creation of the critical size tibiadefect and subsequent Gelrin™ implantation. The observed crack with theempty spaces around the septum (long arrows) is partly a result ofsectioning the paraffin block which contains two different tissues(i.e., bone and fibrous tissue) exhibiting different biomechamicalproperties (i.e., a sectioning artifact). Note the presence of a fibroustissular septum inside the “crack”. Also note the well-formed Haversiansystems (short arrows) representing part of the preexisting cortex and acompacta bone.

FIG. 15 is a photomicrograph of a histological section of rat tibiadepicting a characteristic view of primary bone at 5 weeks followingimplantation of Gelrin™ in the defect site of a critical size defect inthe rat tibia. Note the cartilage undergoing endochondral ossification(in the bottom, left-hand side of the image) as evidenced by theingrowth of vascularized fibrous tissue into the chondroid matrix(denoted by asterisks). The striking feature is the presence of rows ofcuboidal osteoblasts (OST) along the newly formed bone. There is even acluster of osteoblasts within the cartilage itself (long arrow). Thediagnosis of primary bone is based on the presence of chondroid(basophilic) rests within the bone (short arrows). The presence ofenlarged Haversian-like canals with the rows of cuboidal osteoblastsabutting on the bone, indicates the presence of extensive appositionalosteogenesis at this stage following treatment.

FIG. 16 is a photomicrograph of a histological section showing aprimitive—primary bone in a Gelrin™-treated rat critical size tibialdefect at 5 weeks following implantation. Note the abundance of residualbeams of the initially present cartilage (asterisks), which at thisstage has been replaced by osseous tissue. There is a complex network ofchannels. Note that the large Haversian-like canals in the center is atleast partly hypervascularized, and contains a few lipocytes (longarrow). Most of the canal-bone interfaces are covered by cuboidalosteoblasts (small arrows), indicative of ongoing lively osteogenesis.

FIG. 17 is a low-power photomicrograph showing a band of transitioningchondroid tissue in the gap region of a Gelrin™-treated animal. Note theband of cartilage (upper end, indicated by asterisk) in which atransition from fibrous tissue into chondroid tissue is evident(transition occurs from top to bottom, respectively). On both sides ofthis band, many vascularized fibrous tissular projections (short arrows)are invading the cartilage, i.e., initiating endochondral ossification.Although the bone is of osteonal pattern with numerous normal Haversiancanals (long arrows), it is abnormally structured in as much as thereare several large fibro-fatty tissular tracks with rows of cuboidalosteoblasts abutting on the bone (indicated by arrow heads).

FIGS. 18 a-e are phase-contrast microscopic images of vascular smoothmuscle cell cultures inside Gelrin™ Hydrogels. Homogeneously distributedand dispersed smooth muscle cells were cultured for 48 hours insideGelrin™ hydrogels (10-kDa PEG) and the ability of the cells to attachand spread inside the hydrogels was qualitatively assessed by phasecontrast microscopy. FIG. 18 a illustrates the three-dimensionalattachment and spreading of smooth muscle cells inside the pure Gelrin™hydrogels. The addition of varying amounts of free PEG-DA (FIG. 18 b-e)to the Gelrin™ matrix reduced the proteolytic degradability of thehydrogels and makes it more difficult for individual cells to spreadwith in the matrix. FIG. 18 a—pure Gelrin™ matrix; FIG. 18 b—Gelrin™matrix with 0.5% free PEG-DA; FIG. 18 c—Gelrin™ matrix with 1% freePEG-DA; FIG. 18 d—Gelrin™ matrix with 1.5% free PEG-DA; FIG. 18e—Gelrin™ matrix with 2% free PEG-DA. Note that the attachment andspreading of the cells inside the matrix is reduced in the presence ofincreasing concentrations of free PEG-DA (i.e., the cross-linkingmolecule) to the pure Gelrin™ matrix.

FIGS. 19 a-l are photomicrographs of smooth muscle cell clustersembedded within a Gelrin™ substrate. The figure depicts the effect ofthe increasing concentrations of PEG-DA in the Gelrin™ hydrogels onmigration of the cells into the Gelrin™. Gelrin™ hydrogels consisting of6-kDa PEG were prepared using pure Gelrin™ hydrogels (1.75%PEG-fibrinogen; FIGS. 19 a-d), Gelrin™ hydrogels cross-linked with 1%free PEG-DA (FIGS. 19 e-h) or Gelrin™ hydrogels cross-linked with 2%free PEG-DA (FIGS. 19 i-l) and the degree of cell migration from thecellularized tissue mass (dark) and into the Gelrin™ was detected usingphase contrast microscopy following one (FIGS. 19 a, e, and i), two(FIGS. 19 b, f and j), four (FIGS. 19 c, g, and k), or seven (FIGS. 19d, h and i) days in culture. Note the significant cell migration seen inpure Gelrin™ hydrogels and the relatively decreased cell extensionsobserved in Gelrin™ hydrogels which were cross-linked in the presence of2% free PEG-DA.

FIGS. 20 a-c are radiographic images showing new bone formation intreated rats at 5 weeks following the induction of a critical sizedefect. Intermediate degrading hydrogel-treated rats (treatment-2)exhibit extensive new bone formation in the site of the defect asindicated by the formation of a periosteal callus (arrows); the extentof osteoneogenesis ranges from a total bony bridge of the defect (FIG.20 a) to a partially regenerated bone in the gap (FIG. 20 b). All othertreatments are similar to non-treated (control) rats (FIG. 20 c) whichdisclose no new bone formation by radiographic imaging.

FIGS. 21 a-c are photomicrographs of longitudinal sections ofintermediate-degrading hydrogel-treated tibial defects at 5 weeksfollowing the induction of a critical size defect stained withhematoxylin-eosin (H&E). The extent of regenerated bone in thesite-specific defect ranges from partial (FIG. 21 a, FIG. 21 b) to totalbridging (FIG. 21 c) of the defect osteotomies (ost), and highly dependson the erosion pattern of the hydrogel material (Gel). Remnants of thegel give way to regenerated bone (dashed line), having typicallamellar-fibred pattern of mature osseous trabeculae and fatty marrow(FM).

FIGS. 22 a-b are photomicrographs of newly formed subperiosteal andendosteoal bone shown with partially degraded hydrogel in a longitudinalsection of an intermediate-degrading hydrogel-treated rat. FIG. 22 ashows the osseous trabeculae, which conncet with one another and arerimmed by active cuboidal osteoblasts. The intertrabecular spaces areoccupied by a fatty marrow (FM), which extends well into the site of thedefect from the aspect of the medial osteotomy (ost). A cartilaginouscap (arrows) at the medial end of the front of the regenerated bone isseen with islets of hypertrophic chondrocytes. The cap is enclosed by athin layer of perichondrium-like fibrous tissue. Fibro-fatty tissue (FT)is present in between the degraded hydrogel and the regenerated bone.FIG. 22 b is a higher magnification of the same field as in FIG. 22 a.This field displays endochondrol ossification (ECO) in the cartilaginousregion. The sections are stained with hematoxylin and eosin (H&E).

FIGS. 23 a-c are photomicrographs showing that the extent of hydrogeldegradation affects bone healing response at 5 weeks following theinduction of a critical size defect. FIG. 23 a shows that the presenceof the fast-degrading hydrogels (treatment-1) within the site of theimplant results in nonunion, the gap being occupied by fibrousconnective tissue (CT) with a minor mononuclear celled inflammatoryinfiltrate. FIG. 23 b shows the defect site of a rat with an implant ofthe slow degrading hydrogels (treatment-3) filled with the hydrogel,which is enclosed within a fibrous capsule and regenerated bone at theosteotomy (ost) locale. FIG. 23 c shows the intermediate-degradinghydrogel being eroded by granulation tissue and subjacent newly formedbone (NB). The sections are stained with hematoxylin and eosin (H&E).

FIGS. 24 a-b are photomicrographs showing the cellular response to thePEG-fibrinogen Hydrogel Implant. FIG. 24 a illustrates the classicalserpentine granulation tissue at the eroding front of the hydrogel(solid arrows) with adjacent nonspecific chronic inflammatoryinfiltrate, which is primarily composed of lymphocytes (LR) and isaccompanied by newly formed bone (NB). In certain areas of thetissue-material interface the response is limited to only a minorchronic nonspecific inflammatory reaction (dashed arrow). Theeosinophilic hydrogel is lightly stained and shows no cellularinfiltration beyond the eroding borders of the dense matrix (gel). FIG.24 b is a high magnification micrograph showing the pallisadinggranulation tissue; of note is the minor macrophagic reaction (MR).

FIGS. 25 a-c are photographs showing the encapsulation of a singledorsal root ganglion (DRG) into a PEGylated fibrinogen hydrogel. FIG. 25a illustrates that the DRG (arrow) is roughly 0.5 mm in diameter and issituated in the center of a 10 mm diameter PEGylated fibrinogenhydrogel. Radial outgrowth is measured from the outer boundary of theopaque DRG into the transparent hydrogel. The same constructs is shownfrom the top view (FIG. 25 b) and the side view (FIG. 25 c).

FIGS. 26 a-d are photomicrographs showing outgrowth and cellularinvasion characteristics of DRGs in PEGylated fibrinogen 3-D hydrogelconstructs. Phase-contrast micrographs (FIGS. 26 a and 26 b) show thethree-dimensional outgrowth of neurites (arrow) and glial cells(arrowhead) extending from the DRG (D) into the transparent PEGylatedfibrinogen hydrogel construct (P) following two days in culture.Histological sections stained with H&E (FIGS. 26 c and 26 d) of the DRG(dark) in the PEGylated fibrinogen construct (light) show neurite(arrow) and non-neuronal cells (arrowhead) invading the hydrogels afterfour days in culture. Note: high magnification images (FIGS. 26 b and 26d) are expanded regions from the lower magnification micrographs (FIGS.26 a and 26 c); in all images the scale bar=100 μm.

FIGS. 27 a-f are fluorescent microscope images of DRGs encapsulated inPEGylated fibrinogen constructs confirming the presence of both neuritesand schwann cells. Cross sections of DRG constructs were cultured forfour days and fluorescently triple-labeled with βIII-tubulin (neuritemarker, FIGS. 27 a and 27 d), s100 (Schwann cell marker, FIGS. 27 b and27 e), and DAPI counter-stain (nuclei, blue). The merged micrographs(FIGS. 27 c and 27 f) show the three-dimensional invasion of neuritesfrom the DRG into the hydrogel construct, with Schwann cells associatedvery closely with the neurite extensions (scale bar=50 μm).

FIGS. 28 a-d are fluorescent microscope images and graphs showing thatboth free soluble and enmeshed nerve growth factor (FS-NGF and EN-NGF)promote 3-D neurite outgrowth from encapsulated DRGs into the hydrogels.Sections of DRG constructs following four days were immunofluorescentlylabeled for βIII-tubulin (neurites, red), s100 (Schwann cells, green)and DAPI nuclear stain (blue) to characterize the invasion into thehydrogels containing FS-NGF (FIG. 28 a), EN-NGF (FIG. 28 b), or no NGF(NO-NGF; FIG. 28 c). Absence of NGF (NO-NGF) did not encourage outgrowthof neurites but supported moderate outgrowth of Schwann cells.Treatments with free soluble or enmeshed NGF exhibited impressiveoutgrowth of both neurites and Schwann cells (scale bar=50 μm). FIG. 28d is a line graph illustrating that the average neurite extension length(±standard error) in free soluble NGF treatments (FS-NGF) and enmeshedNGF treatments (EN-NGF) is not significantly different between the twotreatments (P>0.35, n=6).

FIGS. 29 a-u are phase contrast micrographs and graphs showing thatPEG-fibrinogen composition controls neurite invasion and outgrowth. Thehydrogel composition is varied using different amounts of PEG andfibrinogen during assembly, including 30:1, 60:1, 120:1, and 180:1(PEG:Fibrinogen). FIGS. 29 a-p are phase contrast micrographs showingthe profound impact of additional PEG on the 3-D outgrowth morphologyfrom the DRG following four days (scale bar=200 μm). FIGS. 29 q-t arehigh magnification images showing the relative outgrowth of neurites andglial cells into hydrogels having different compositions (scale bar=200μm). FIG. 29 u is a line graph showing that the average neuriteextension length (±standard error) in each treatment, as measureddirectly from the images, shows no significant difference between the30:1 and 60:1 treatments (P>0.50, n=9), and a significant impediment tooutgrowth in the 120:1 and 180:1 treatments (p<0.01, n=9).

FIGS. 30 a-c are phase contrast micrographs showing DRG outgrowth intohydrogels made from PEG-DA and PEG-fibrinogen. Constructs were preparedwith 10% PEG-DA gels (w/v) without fibrinogen (FIG. 30 a) and comparedwith PEG-fibrinogen constructs (FIG. 30 b). Neurite extensions werebarely visible following three days in PEG-DA construct compared toextensive invasion seen in PEGylated fibrinogen constructs after threedays. Neuronal invasion into PEGylated fibrinogen hydrogels waseliminated in the absence of NGF (FIG. 30 c), although other cells typeswere observed in the hydrogel following three days in culture. In allimages the scale bar=200 μm.

DESCRIPTION OF THE PREFERRED EMBODIMENTS

The present invention is of a matrix composed of a naturally-occurringprotein backbone cross-linked by polyethylene glycol (PEG) which can beused in tissue regeneration applications. Specifically, the presentinvention is of a PEGylated fibrinogen scaffold which can be used totreat disorders characterized by tissue damage or loss usingbiodegradable scaffolds.

The principles and operation of the method of generating a biodegradablescaffold according to the present invention may be better understoodwith reference to the drawings and accompanying descriptions.

Before explaining at least one embodiment of the invention in detail, itis to be understood that the invention is not limited in its applicationto the details set forth in the following description or exemplified bythe Examples. The invention is capable of other embodiments or of beingpracticed or carried out in various ways. Also, it is to be understoodthat the phraseology and terminology employed herein is for the purposeof description and should not be regarded as limiting.

Tissue engineering approaches utilize basic scaffolds which mimic thenatural structure of the tissues they replace and provide temporaryfunctional support for cells seeded thereon (Griffith L G, 2002).Scaffolds can be fabricated from either biological materials orsynthetic polymers. While synthetic polymers [e.g., polyethylene glycol(PEG) and Polymethyl methacrylate (PMMA)], provide precise control overthe scaffold material and its mechanical properties, such polymers aredevoid of biofunctional properties which enable cell attachment andspreading (Drury and Mooney, 2003), and are therefore unsuitable forlong-term tissue culture or in vivo tissue regeneration.

Naturally occurring scaffolds such as collagen and fibrin providebio-functional signals and exhibit various cellular interactions.However, since such biological materials exhibit multiple inherentsignals (e.g., regulation of cell adhesion, proliferation, cellularphenotype, matrix production and enzyme activity), their use asscaffolds for tissue regeneration often results in abnormal regulationof cellular events (Hubbell, 2003). In addition, followingreconstitution, the strength of such scaffolds is lower than thatcharacterizing the natural biological material (e.g., collagen).

To overcome such limitations “hybrid” scaffolds were constructed bygrafting biodegradable elements into synthetic backbones. Thus,synthetic PEG backbone was cross-linked with short oligopeptidescontaining enzymatic substrates, RGD and/or VEGF sequences [Lutolf et al(2003); Gobin and West (2002); Seliktar et al; 2004, Zisch A H, et al,2003; FASEB J. 17: 2260-2. Epub 2003 Oct. 16], or withgenetically-engineered protein-like precursors of 100 amino acids(Halstenberg et al. 2002; Biomacromolecules, 3: 710-23). However, suchscaffolds, in which PEG was the major component, failed to providesufficient bio-feedback and/or long-term cellular responses which areessential for tissue regeneration and phenotypic stability. Animprovement of such scaffold is characterized by the addition of othergrowth factors to the premade scaffold, however in this case the growthfactor does not form the scaffold but rather is an auxiliary componentthereof [Lutolf M P, Weber F E, Schmoekel H G, Schense J C, Kohler T,Muller R, Hubbell J A. Repair of bone defects using synthetic mimeticsof collagenous extracellular matrices. Nat Biotechnol. 2003May;21(5):513-8].

Other attempts made by Fortier and co-workers consisted of a single stepscaffold generation process, which resulted in a scaffold formed fromnaturally occurring proteins which is capable of mediating biologicalcues. However, in the absence of further purification steps, such ascaffold included toxic unreacted amine-reactive di-functional PEG,which disqualified its use in the clinic [Jean-Francois J, D'Urso E M,Fortier G. Immobilization of L-asparaginase into a biocompatiblepoly(ethylene glycol)-albumin hydrogel: evaluation of performance invivo. Biotechnol Appl Biochem. 1997 Dec.;26 (Pt 3):203-12; Jean-FrancoisJ, Fortier G. Immobilization of L-asparaginase into a biocompatiblepoly(ethylene glycol)-albumin hydrogel: I: Preparation and in vitrocharacterization. Biotechnol Appl Biochem. 1996 Jun.;23 (Pt 3):221-6;Gayet J C, Fortier G. Drug release from new bioartificial hydrogel.Artif Cells Blood Substit Immobil Biotechnol. 1995;23(5):605-11; D'UrsoE M, Jean-Francois J, Doillon C J, Fortier G. Poly(ethyleneglycol)-serum albumin hydrogel as matrix for enzyme inimobilization:biomedical applications. Artif Cells Blood Substit Immobil Biotechnol.1995;23(5):587-95].

While reducing the present invention to practice, the present inventorshave uncovered that biosynthetic hybrid scaffolds composed of afibrinogen backbone with functional polyethylene glycol (PEG) sidechains are excellent, biodegradable scaffolds and that such scaffoldscan be used for tissue regeneration applications.

As is shown in the Examples section which follows, the present inventorsgenerated PEG-fibrinogen precursor molecules which were further used toform hydrogel scaffolds. The PEG-fibrinogen scaffolds of the presentinvention exhibit material properties and biodegradability which aresuperior to any known prior art scaffolds; they also exhibit highflexibility and a controllable elastic modulus (FIGS. 3 a-b), efficientbiodegradability (FIGS. 5 a-b), improved biofunctionality and supportfor cell spreading and extension (FIGS. 6 a-e, 7 a-b, 18 a-e andExamples 2 and 4). As shown in FIGS. 9-17 and FIGS. 20 a-c and FIGS. 23a-c in Examples 3 and 5 of the Examples section which follows, thePEG-fibrinogen scaffolds of the present invention were capable of invivo bone regeneration in rats exposed to critical size tibia defect.

Moreover, as illustrated in FIGS. 26-30 in Example 6, the PEG-fibrinogenscaffolds of the present invention were also capable of ex vivo nerveregeneration.

In sharp contrast to the above-described scaffolds consisting ofmulti-functional PEG network cross-linked with reactive syntheticoligopeptide, the present invention employs a natural protein PEGylatedwith di-functional or multi-functional PEG to form the backboneprecursor molecule. The naturally occuring PEGylated protein of thepresent invention is both the degradation substrate and the mainconstituent of the biological signaling of the matrix. Thebiodegradation is controlled by the composition of PEG and protein: thedegree of PEGylation, the relative amount of PEG and protein, and thelength of the PEG molecules. The protein backbone also serves as themain signaling constituent of the scaffold, compared to the syntheticsystem where factors (growth factors) and other biofunctional syntheticoligopeptides (RGD) can be tethered to the multi-functional PEG networkthrough an additional reaction in order to create further biologicalsignals required for tissue healing. Because the biological activity isinherent to the protein backbone of the present invention, the samefragments of the denatured PEGylated protein which may contain manyinductive properties that are relevant for physiological injury responsecan contribute to the observed healing of the defect as the matrix isdegraded.

Thus, according to one aspect of the present invention there is provideda composition-of-matter comprising a naturally occurring protein or abioactive fragment thereof and at least two synthetic polymerscovalently attached thereto, each of said at least two syntheticpolymers having a functional group being capable of attaching to saidnaturally occurring protein or said bioactive fragment thereof so as toform a scaffold.

As used herein the phrase “scaffold” refers to a two-dimensional or athree-dimensional supporting framework. The scaffold of the presentinvention is composed of units (interchangeably referred to herein as“precursor molecules”) which are directly or indirectly (e.g., vialinker) attachable therebetween. Such precursor molecule can be forexample, PEGylated fibrinogen (see Example 1 of the Examples sectionwhich follows), PEGylated collagen, PEGylated fibronectin and the like.By controlling cross-linking, the scaffold of the present invention canform two- or three-dimensional structure at any size, structure orporosity. The scaffold of the present invention can be embedded within,or formed around, another scaffold or gel or it can be linked toadditional materials to form a hybrid or coated scaffold.

Preferably, the scaffold of the present invention can be used to supportcell growth, attachment, spreading, and thus facilitate cell growth,tissue regeneration and/or tissue repair.

The term “polymer” refers to a plurality of repeating units which form anew molecular structure. The phrase “synthetic polymer” refers to anypolymer which is made of a synthetic material, i.e., a non-natural,non-cellular material. Non-limiting examples for synthetic polymerswhich can be used along with the present invention include polyethyleneglycol (PEG) (average Mw. 200; P3015, SIGMA),Hydroxyapatite/polycaprolactone (HA/PLC) [Choi, D., et al., 2004,Materials Research Bulletin, 39: 417-432; Azevedo M C, et al., 2003, J.Mater Sci. Mater. Med. 14(2): 103-7], polyglycolic acid (PGA) [NakamuraT, et al., 2004, Brain Res. 1027(1-2): 18-29], Poly-L-lactic acid (PLLA)[Ma Z, et al., 2005, Biomaterials. 26(11): 1253-9], Polymethylmethacrylate (PMMA) [average Mw 93,000, Aldrich Cat. # 370037; Li C, etal., 2004, J. Mater. Sci. Mater. Med. 15(1): 85-9], polyhydroxyalkanoate(PHA) [Zinn M, et al., 2001, Adv. Drug Deliv. Rev. 53(1): 5-21; SudeshK., 2004, Med. J. Malaysia. 59 Suppl B: 55-6], poly-4-hydroxybutyrate(P4HB) [Dvorin E L et al., 2003, Tissue Eng. 9(3): 487-93],polypropylene fumarate (PPF) [Dean D, et al., 2003, Tissue Eng. 9(3):495-504; He S, et al., 2000, Biomaterials, 21(23): 2389-94],polyethylene glycol-dimethacrylate (PEG-DMA) [Oral E and Peppas N A J,2004, Biomed. Mater. Res. 68A(3): 439-47], beta-tricalcium phosphate(beta-TCP) [Dong J, et al., 2002, Biomaterials, 23(23): 4493-502], andnonbiodegradable polytetrafluoroethylene (PTFE) [Jernigan T W, et al.,2004. Ann. Surg. 239(5): 733-8; discussion 738-40].

The synthetic polymers of this aspect of the present invention may beidentical or different. As mentioned herein-above, each of the syntheticpolymers in the composition comprises a functional group which iscapable of forming a direct or indirent bond with the naturallyoccurring protein such as to a side chain thereof or to an end group.Such a functional group may comprise an amine a thiol and the like.

According to presently preferred embodiments of the present inventionthe synthetic polymer used by the present invention is PEG. The PEGmolecule used by the present invention can be linearized or branched(i.e., 2-arm, 4-arm, and 8-arm PEG) and can be of any molecular weight,e.g., 4 kDa, 6 kDa and 20 kDa for linearized or 2-arm PEG, 14 kDa and 20kDa for 4-arm PEG, and 14 kDa and 20 kDa for 8-arm PEG and combinationthereof.

As is shown in FIGS. 1 a-b and Example 1 of the Examples section whichfollows the OH-termini of the PEG molecule can be reacted with achemical group such as acrylate (Ac) or vinylsulfone (VS) which turn thePEG molecule into a functionalized PEG, i.e., PEG-Ac or PEG-VS.Preferably, the PEG molecule used by the present invention is PEG-Ac.

Methods of preparing functionalized PEG molecules are known in the arts.For example, PEG-VS can be prepared under argon by reacting adichloromethane (DCM) solution of the PEG-OH with NaH and then withdi-vinylsulfone (molar ratios: OH 1: NaH 5: divinyl sulfone 50, at 0.2gram PEG/mL DCM). PEG-Ac is made under argon by reacting a DCM solutionof the PEG-OH with acryloyl chloride and triethylamine (molar ratios: OH1: acryloyl chloride 1.5: triethylamine 2, at 0.2 gram PEG/mL DCM),essentially as described in Example 1 of the Examples section whichfollows.

It will be appreciated that such chemical groups can be attached tolinearized, 2-arm, 4-arm, or 8-arm PEG molecules.

Preferably, the PEG-Ac used by the present invention is PEG-DA, 4-armstar PEG multi-Acrylate and/or 8-arm star PEG multi-Acrylate.

As is shown in FIGS. 2 a-b and Example 1 of the Examples section whichfollows the present inventors used 4-kDa, 6-kDa and 20-kDa isoforms ofPEG-diacrylate (PEG-DA) to prepare functionalized PEG molecules.

According to a presently known preferred embodiment of the presentinvention, a 6-14 kDa PEG-diacrylate is used.

The phrase “naturally occurring protein or a portion thereof” as usedherein refers to any peptide, polypeptide or protein which exists innature, such as in eukaryotic and/or prokaryotic organisms, cells,cellular material, non-cellular material and the like. Such a protein ora portion thereof may have a known or unknown structure, function, ormolecular properties. For example, the protein of the present inventioncan be a cell signaling protein, an extracellular matrix protein, a celladhesion protein, a growth factor, a protease, and a protease substrate.

Examples for cell signaling proteins include, but are not limited to,p38 mitogen-activated protein kinase (GenBank Accession No.NP_(—)002736), nuclear factor kappaB (GenBank Accession No.NP_(—)003989), Raf kinase inhibitor protein (RKIP) (GenBank AccessionNo. XP_(—)497846), Raf-1 (GenBank Accession No. NP_(—)002871), MEK(GenBank Accession No. NP_(—)002746), Protein kinase C (PKC) (GenBankAccession No. NP_(—)002728), phosphoinositide-3-kinase gamma (GenBankAccession No. NP_(—)002640), receptor tyrosine kinases [e.g., insulinreceptor (GenBank Accession No. NP_(—)000199)], heterotrimericG-proteins [e.g., Galpha(i) (GenBank Accession No. NP_(—)002060),Galpha(s) NP_(—)000507 and Galpha(q) (GenBank Accession No.NP_(—)002063)], Caveolin-3 (GenBank Accession No. NP_(—)001225), 14-3-3proteins (GenBank Accession No. NP_(—)003397).

Examples for extracellular matrix proteins include, but are not limitedto, fibrinogen [α-chain—GenBank Accession No. NP_(—)068657 (SEQ IDNO:4); β-chain—GenBank Accession No. P02675 (SEQ ID NO:5);γ-chain—GenBank Accession No. P02679 (SEQ ID NO:6)], Collagen (GenBankAccession No. NP_(—)000079), fibronectin (GenBank Accession No.NP_(—)002017), vimentin (GenBank Accession No. NP_(—)003371),microtubule-associated protein lb (GenBank Accession No. NP_(—)005900)(Theodosis D T. 2002; Front Neuroendocrinol. 23: 101-35), Neuriteoutgrowth factor (NOF) (GenBank Accession No. P21741) (Tsukamoto Y, etal., 2001; Histol. Histopathol. 16: 563-71), bacterial cellulose (BC)(GenBank Accession No. NP_(—)625477), laminin (GenBank Accession No.NP_(—)000218) and gelatin [Zhang Y., et al., 2004; J Biomed Mater Res.2004 Sep. 22; Epub ahead of print].

Examples for cell adhesion proteins include, but are not limited to,integrin (GenBank Accession No. NP_(—)002202) (Stefanidakis M, et al.,2003; J Biol. Chem. 278: 34674-84), intercellular adhesion molecule(ICAM) 1 (GenBank Accession No. NP_(—)000192) (van de Stolpe A and vander Saag P T. 1996; J. Mol. Med. 74: 13-33), N-CAM GenBank Accession No.NP_(—)000606), cadherin (GenBank Accession No. NP_(—)004351), tenascin(GenBank Accession No. NP_(—)061978) (Joshi P, et al., 1993; J. CellSci. 106: 389-400), gicerin (GenBank Accession No. NP_(—)006491), andnerve injury induced protein 2 (ninjurin2) (GenBank Accession No.NP_(—)067606) (Araki T and Milbrandt J. 2000; J. Neurosci. 20: 187-95).

Examples of growth factors include, but are not limited to, EpidermalGrowth Factor (GenBank Accession No. NP_(—)001954), transforming growthfactor-beta (GenBank Accession No. NP_(—)000651), fibroblast growthfactor-acidic (GenBank Accession No. NP_(—)000791), fibroblast growthfactor-basic (GenBank Accession No. NP_(—)001997), erythropoietin(GenBank Accession No. NP_(—)000790), thrombopoietin (GenBank AccessionNo. NP_(—)000451), hepatocyte growth factor (GenBank Accession No.NP_(—)000592), insulin-like growth factor-I (GenBank Accession No.NP_(—)000609), insulin-like growth factor-II (GenBank Accession No.NP_(—)000603), Interferon-gamma (GenBank Accession No. NP_(—)000610),and platelet-derived growth factor (GenBank Accession No. NP_(—)079484).

Examples for protease proteins include, but are not limited to, pepsin(GenBank Accession No. NP_(—)055039), low specificity chymotrypsin, highspecificity chymotrypsin, trypsin (GenBank Accession No. NP_(—)002760),carboxypeptidases (GenBank Accession No. NP_(—)001859), aminopeptidases(GenBank Accession No. NP_(—)001141), proline-endopeptidase (GenBankAccession No. NP_(—)002717), Staphylococcus aureus V8 protease (GenBankAccession No. NP_(—)374168), Proteinase K (PK) (GenBank Accession No.P06873), aspartic protease (GenBank Accession No. NP_(—)004842), serineproteases (GenBank Accession No. NP_(—)624302), metalloproteases(GenBank Accession No. NP 787047), ADAMTS17 (GenBank Accession No.NP_(—)620688), tryptase-gamma (GenBank Accession No. NP_(—)036599),matriptase-2 (GenBank Accession No. NP_(—)694564).

Protease substrate proteins include the peptide or peptide sequencesbeing the target of the protease protein. For example, Lysine andArginine are the target for trypsin; Tyrosine, Phenylalanine andTryptophan are the target for chymotrypsin.

Such naturally occurring proteins can be obtained from any knownsupplier of molecular biology reagents such as Sigma-Aldrich Corp., StLouis, Mo., USA and Invitrogen Carlsbad Calif.

As is mentioned hereinabove, portions (fragments e.g., bioactivefragments capable of mediating biological funciton) of such proteins canbe also used to generate the composition of the present invention. Sucha portion usually includes sufficient biodegradability potential, i.e.,protease substrates and/or protease targets, as well as sufficient cellsignaling and/or cell adhesion motives. For example, the humanfibrinogen protein contains two RGD adhesion sites at amino acids114-116 and 591-593 of the α-chain as set forth in SEQ ID NO:4 (GenBankAccession No. NP_(—)068657), as well as a protease cleavage site atamino acids 44-45 of the β-chain as set forth in SEQ ID NO:5 (GenBankAccession No. P02675).

The above described composition of matter may be cross-linked to form ascaffold.

Thus, according to another aspect of the present invention there isprovided a scaffold comprising a plurality of protein moleculescovalently attached therebetween so as to form the scaffold via asynthetic polymer (such as described hereinabove) having a first partand a second part covalently connected therebetween via a chemicalmoiety being chemically distinct from a repeating unit of said polymer.

Examples of such chemical moieties include, but are not limited to,aldehyes, acetale, tosyl, tresyl, dichlorotriazine, epoxide, carboxylic,succinimidyle succinate, succinimidyl ester, p-nitrophenyl carbonate,benzotriazolyl carbonate, 2,3,5-trichlorophenyl carbonate, succinimidylecarbonate, pyridildisulphide, maleimide, vinylsulfone, and iodoacetamide

Since the scaffold of the present invention is formed from precursormolecules (i.e., the above-described composition of matter) it is devoidof uncojugated forms of the synthetic polymer (such as below 20%,preferably below 15%, preferably below 10% and preferably below 5%),rendering it biologically safe, and particularly useful for clinicalapplications.

Since the scaffold of the present invention is composed of a naturallyoccurring protein such a scaffold can be configured susceptible todegradation by biological material such as enzymes, i.e., biodegradable.

As used herein, the phrase “biodegradable” refers to capable of beingdegraded (i.e., broken down) by biological proteases or biomolecules.Biodegradability depends on the availability of degradation substrates(i.e., biological materials or portion thereof), the presence ofbiodegrading materials (e.g., microorganisms, enzymes, proteins) and theavailability of oxygen (for aerobic organisms, microorganisms orportions thereof), carbon dioxide (for anaerobic organisms,microorganisms or portions thereof) and/or other nutrients. In addition,biodegradability of a material, such as the scaffold of the presentinvention, also depends on the material structure and/or mechanicalproperties, i.e., the porosity, flexibility, viscosity, cross-linkdensity, hydrophobicity/hydrophilicity, and elasticity which may affectpassage and availability of gasses and nutrients, as well as cellattachment and spreading. Examples of biodegradable materials include,but are not limited to, reconstituted collagen gels, fibrin glues, andhyaluronic acid scaffolds.

Thus, the scaffold of the present invention is fabricated by crosslinking multiple copies of a precursor molecule (interchangeabledescribed as the composition-of-matter of the present invention) whichis composed of a synthetic polymer attached to a naturally occurringprotein or a portion thereof.

For example, as is shown in FIGS. 1 a-b, 2 a-b and Example 1 of theExamples section which follows, the present inventors fabricated aPEGylated fibrinogen precursor molecule by denaturing fibrinogenmolecules and reacting them with functionalized PEG-diacrylates.

Thus, according to preferred embodiments of the present invention, thepolymer-protein precursor molecule is made of PEG and fibrinogen.

The molar ratio between the synthetic polymer (e.g., PEG) and thenaturally occurring protein (e.g., fibrinogen) of the present inventionmay affect the pore size, strength, flexibility and elasticity of thescaffold of the present invention. Thus, excess of the synthetic polymerwould lead to binding of the polymer functional groups (e.g., PEG-DA) toall potential binding sites on the naturally occurring protein and wouldresult in smaller pore size, more cross-linking sites, higher strength,less flexibility and increased rigidity. On the other hand, binding ofonly two molecules of the synthetic polymer to each molecule of theprotein (i.e., a 2:1 molar ratio) would result in large pore size, fewercross-linking sites, lower strength, higher flexibility and increasedelasticity. It will be appreciated that the molar ratio between thesynthetic polymer and the protein can also affect the biodegradabilityof the scaffold. Thus, a higher molar ratio (i.e., excess of polymer) isexpected to result in less biodegradability due to potential masking ofprotein degradation sites. Those of skills in the art are capable ofadjusting the molar ratio between the synthetic polymer and the proteinto obtain the desired scaffold with the optimal physical and biologicalcharacteristics.

For example, since each fibrinogen molecule includes 29-31 potentialsites which can bind to PEG, the PEG-fibrinogen precursor molecule canbe prepared using a wide range of molar ratios. Preferably, the molarratio used by the present invention is 2-400 (PEG) to 1 (fibrinogen),more preferably, the molar ratio is 30-300 (PEG) to 1 (fibrinogen), morepreferably, the molar ratio is 100-200 (PEG) to 1 (fibrinogen), mostpreferably, the molar ratio is 130-160 (PEG) to 1 (fibrinogen). As isshown in Example 1 of the Examples section which follows, one preferablemolar ratio between PEG-DA and fibrinogen is 145 (PEG) to 1(fibrinogen). In the case where the molar ratio is greater than 29-31(PEG) to 1 (fibrinogen), some of the PEG can be indirectly bound to thefibrinogen through fibrinogen-bound PEG molecules.

The fibrinogen used by the present invention can be whole denaturedfibrinogen (i.e., un-cleaved) or fragmented fibrinogen, which can beobtained using, for example, CNBr cleavage (see Example 1 of theExamples section which follows).

Fibrinogen can be readily purified from human blood plasma usingstandard protein purification techniques. Purified components may besubject to anti-viral treatments. Heat-treatment and solvent/detergenttreatments are both commonly used in the production of fibrinogen.Fibrinogen used in accordance with the present invention is preferablypure though other components may be present. Thus products may alsocontain tranexamic acid, aprotinin or factor XIII. Fibrinogen iscommercially available from Baxter and Omrix.

As is mentioned before, the scaffold of the present invention is formedby cross-linking the polymer-protein precursor molecules of the presentinvention. Such cross-linking can be performed in vitro, ex vivo and/orin vivo.

Cross-linking according to this aspect of the present invention isperformed by subjecting the precursor molecules to a free-radicalpolymerization reaction (i.e., a cross-linking reaction). Methods ofcross-linking polymers are known in the art, and include for example,cross-linking via photoinitiation (in the presence of an appropriatelight, e.g., 365 nm), chemical cross-linking [in the presence of afree-radical donor] and/or heating [at the appropriate temperatures.Preferably, cross-linking according to the present invention is effectedby photoinitiation.

Photoinitiation can take place using a photoinitiation agent (i.e.,photoinitiator) such as bis(2,4,6-trimethylbenzoyl) phenylphosphineoxide (BAPO) (Fisher J P et al., 2001; J. Biomater. Sci. Polym. Ed. 12:673-87), 2,2-dimethoxy-2-phenylacetophenone (DMPA) (Witte R P et al.,2004; J. Biomed. Mater. Res. 71A(3): 508-18), camphorquinone (CQ),1-phenyl-1,2-propanedione (PPD) (Park Y J et al., 1999, Dent. Mater.15(2): 120-7; Gamez E, et al., 2003, Cell Transplant. 12(5): 481-90),the organometallic complex Cp′Pt(CH(3))(3) (Cp′=eta(5)-C(5)H(4)CH(3))(Jakubek V, and Lees A J, 2004; Inorg. Chem. 43(22): 6869-71),2-hydroxy-1-[4-(hydroxyethoxy)phenyl]-2-methyl-1-propanone (Irgacure2959) (Williams C G, et al., 2005; Biomaterials. 26(11): 1211-8),dimethylaminoethyl methacrylate (DMAEMA) (Priyawan R, et al., 1997; J.Mater. Sci. Mater. Med. 8(7): 461-4), 2,2-dimethoxy-2-phenylacetophenone(Lee Y M et al., 1997; J. Mater. Sci. Mater. Med. 8(9): 537-41),benzophenone (BP) (Wang Y and Yang W. 2004; Langmuir. 20(15): 6225-31),flavin (Sun G, and Anderson V E. 2004; Electrophoresis, 25(7-8):959-65).

The photoinitiation reaction can be performed using a variety ofwave-lengths including UV (190-365 nm) wavelengths, and visible light(400-1100 nm) and at various light intensities. It will be appreciatedthat for ex vivo or in vivo applications, the photoinitiator andwavelengths used are preferably non-toxic and/or non-hazardous.

For example, as is shown in Example 1 of the Examples section whichfollows, the PEG-fibrinogen precursor molecule was cross-linked byphotoinitiation in the presence of Igracure™2959 and a non-toxic UVlight illumination (e.g., 5 minutes at 365 nm wavelength, 4-5 mWatts/cm²intensity).

It will be appreciated that although the polymer-protein precursormolecules of the present invention (e.g., PEGylated fibrinogen) arecapable of being cross-linked without the addition of a cross-linkingmolecule, cross-linking according to the present invention can alsoutilize a molecule capable of cross-linking the polymer-proteinprecursors. Such cross-linking molecules can be for examples, PEG,PEG-DA, PEG multi-Acrylate, and/or PEG-VS.

For example, as is shown in Example 1 of the Examples section whichfollows, a functionalized PEG molecule (e.g., PEG-DA) was used toenhance the cross-linking reaction.

The concentration of cross-linking molecules (e.g., PEG-DA) can affectthe scaffold strength, flexibility, elasticity and biodegradability anddetermination of such a concentration depends on the scaffoldapplication and is within the capabilities of those skilled in the arts.For example, excess of a cross-linking molecule is expected to result insmaller pores, more cross-linking sites, and higher scaffold strengthand less flexibility. On the other hand, as is shown in FIGS. 18 a-e, 19a-1 and Example 4 of the Examples section which follows, excess ofPEG-DA (i.e., the cross-linking molecule of the present invention)resulted in reduced scaffold biodegradability as indicated by thereduced cell attachment and/or spreading. It will be appreciated thatreduced biodegradability is probably a result of masking or modifyingthe protein binding sites or signals which are necessary for proteindegradation.

According to a preferred embodiment of this aspect of the presentinvention, an excess of 3% PEG-DA is added in order to promote boneformation. The present inventors have shown that this concentration ofexcess PEG-DA provides the optimal biodegradability characteristics forthe hydrogel regarding bone formation (FIGS. 20 a-c and FIGS. 23 a-c).Without being bound to theory, it is believed that a sustained releaseof fragments of PEGylated fibrinogen from the hydrogel synchronized withdegradability is necessary to fully capitalize on the osteoinductiveproperties of the hydrogel of the present invention.

According to preferred embodiments of the present invention,cross-linking is effected such that the polymer-protein precursors ofthe present invention are solubilized in a water-based solution and suchsolutions are further subjected to cross-linking (e.g., usingphotoinitiation) to form a hydrogel scaffold.

As is shown in Example 1 of the Examples section which follows, aPEG-fibrinogen hydrogel was formed by mixing the PEGylated fibrinogenprecursor molecules with the photoinitiation agent (Igracure™2959) inthe presence or absence of PEG-DA and exposing such a mixture to UVlight. Briefly, the PEGylated fibrinogen precursors were solubilized in1-ml of 50 mM PBS, pH 7.4 and 25° C. to achieve a final concentration of10, 15, or 20% polymer-protein (w/v). The precursor solution alsocontained a PEG-DA cross-linking constituent at a molar ratio of 1:2PEG-DA to functional groups on the PEGylated fibrinogen. The precursorsolution was mixed with 10 μl of Igracure™2959 photoinitiator solution(Ciba Specialty Chemicals, Tarrytown, N.Y.) in 70% ethanol (100 mg/ml)and centrifuged for 5 min at 14,000 RPM. The solution was then placedinto Teflon tubes (5-mm diameter and 20-mm long) and polymerized underUV light (365 nm, 4-5 mW/cm²) for 15 minutes according to publishedprotocols (Lum L Y et al., 2003).

According to preferred embodiments of the present invention the hydrogelcan be generated from PEGylated whole fibrinogen or PEGylated fragmentedfibrinogen. Generally, the molecular weight and length of the graftedPEG affects the degree of solubility of the PEGylated protein, i.e.,higher length and/or molecular weight of PEG results in increasedsolubility of PEGylated protein. It will be appreciated that solubilityof the PEGylated protein is also affected by the presence of whole orcleaved fibrinogen. Preferably, the concentration of the precursormolecules in the hydrogel is between 0.5 to 35%, more preferably, whenPEGylated whole fibrinogen is used, the concentration of the precursormolecules in the hydrogel is between 0.5 to 5% (depending on the MW andlength of the grafted PEG used to PEGylate the protein) and whenPEGylated fragmented fibrinogen is used, the concentration of theprecursor molecules in the hydrogel is between 5-35% (depending on theMW and length of PEG used to PEGylate the protein).

The PEG-fibrinogen hydrogels of the present invention exhibited asuperior flexibility over the prior art PEG-PEG hydrogels. For example,while in the PEG-PEG hydrogel (using 4 kDa PEG) a strain of 30% wasachieved by employing a stress of 90 kPa, in the PEG-fibrinogen hydrogel(using 4 kDa PEG) a similar strain was achieved by employing only 4 kPa(FIGS. 3 a-b and FIGS. 4 a-c).

Preferably, the modulus of elasticity of the hydrogels made fromPEGylated whole fibrinogen is in a range of 0.02-0.11 kPa for 10-20%polymer, and the modulus of elasticity of the hydrogels made fromPEGylated fragmented fibrinogen is in a range of 0.01-0.07 kPa for10-20% polymer.

In addition, the hydrogel scaffolds of the present invention exhibithigh biodegradability as compared with prior art hydrogel scaffolds[e.g., hydrogels made with an oligopeptide cross-linker containing aprotease substrate (Seliktar et al 2004)]. For example, as is shown inFIGS. 5 a-b and Example 1 of the Examples section which follows, asignificant degradation of 45-70% or 35-85% of the PEG-fibrinogenhydrogels was observed following 30 minutes incubation in the presenceof 0.05 mg/ml trypsin or 0.5 mg/ml collagenase, respectively. Moreover,as is further shown in FIG. 5 b, higher concentrations of trypsin (i.e.,1-2 mg/ml) resulted in a complete hydrogel degradation following 30minutes of incubation.

Thus, the biodegradability of the hydrogel scaffold of the presentinvention can be determined by subjecting such hydrogels to enzymaticdegradation using proteases such as plasmin, trypsin, collagenase,chemotrypsin and the like.

It will be appreciated that the biodegradability and biofunctionality ofthe scaffold hydrogel of the present invention can be further increasedby loading with a pharmaceutical agent of interest.

Thus, for example, compositions of the present invention may enclosecomponents which are nonreactive to the hydrogel. Examples of suchnonreactive components may include drugs such as disinfectants,chemotherapeutics, antimicrobial agents, antiviral agents, hemostatics,antiphlogistics, anesthetics, analgesics, or nutritional supplements;biopolymers such as peptides, plasma derivative proteins, enzymes ormixtures thereof. In other words, components nonreactive to the hydrogelmay be combined with the composition for hydrogel to providestabilization or protection of these components. Such combinedcomposition may be prepared, for example, by dissolving or suspendingthe nonreactive components in the aqueous medium to be used for gelationbefore effecting the gelation. Methods of loading hydrogels withpharmaceuticals are well known in the art [see for example, druginclusion in bovine serum albumin hydrogels desrcribed in Gayet andFortier (1995) Art. Cells. Blood Subs. And Immob. Biotech. 23(5),605-611].

Alternatively or additionally, the biodegradability and biofunctionalityof the scaffold hydrogel of the present invention can be furtherincreased by attaching or impregnating a protein such as a cellsignaling protein, or a growth factor (e.g. a nerve growth factor asdescribed in Example 6 hereinbelow) to the hydrogel of the presentinvention. Attaching such proteins to the hydrogel scaffold of thepresent invention is preferably employed by covalent immobilization of aPEGylated protein to the PEG hydrogel network during cross-linking(Seliktar et al 2004, JBMR). The immobilization of such factor isaccomplished by directly reacting functionalized PEG to an unreactedthiol present on a cysteine residue of the protein sequence.Impregnation of the hydrogel with growth factors can be performed bydehydrating the scaffold and then immersing the hydrogels in a solutioncontaining the growth factors and gently shaking such hydrogels for afew hours until the growth factors penetrate the scaffold during thehydration process. Likewise, the hydrogel can be impregnated with growthfactor by incubation in factor-containing solution overnight until thegrowth factor diffuses into the polymeric network of the scaffold byslow, passive diffusion. The latter is influenced by the degree ofcross-linking, the porosity of the scaffold, and the structuralproperties described hereinabove.

Apart from being inexpensive to produce, the scaffold of the presentinvention is highly reproducible, flexible (can be stressed or stretchedeasily), exhibit a controllable structural properties, and amenable tocontrollable biodegradation; characteristics which make it highlysuitable for in vivo or ex vivo regeneration of tissues such as bone,nerve, cartilage, heart muscle, skin tissue, blood vessels, and othertissues (soft and hard) in the body. For example, such a scaffoldhydrogel can be easily placed into gaps within a tissue or an organ,following which it can fill the void and initiate the process ofregeneration as the scaffold degrades away.

Indeed, as is shown in FIGS. 8-17, FIGS. 20 a-c and FIGS. 23 a-c inExamples 3 and 5 of the Examples section which follows, implantation ofthe scaffold of the present invention in a critical size rat tibiadefect resulted in new bone formation.

Thus, according to another aspect of the present invention there isprovided a method of inducing in vivo formation of a tissue.

The phrase “in vivo” refers to within a living organism such as a plantor an animal, preferably in mammals, preferably, in human subjects.

The method is effected by implanting the scaffold of the presentinvention in a subject to thereby induce the formation of the tissue.

As used herein, the term “subject” refers to a vertebrate, preferably amammal, more preferably a human being (male or female) at any age.

The scaffold of the present invention can be implanted in the subjectusing a surgical tool such as a scalpel, spoon, spatula, or othersurgical device.

It will be appreciated that in vivo formation of a tissue can be alsoachieved by administering the scaffold precursor molecules to thesubject and further cross-linking the precursor molecules in vivo.

Thus, according to another aspect of the present invention there isprovided a method of inducing in vivo formation of a tissue. The methodis effected by administering to a subject in need thereof a compositioncomposed of a synthetic polymer attached to a naturally occurringprotein or a portion thereof, the composition capable of forming ascaffold within the subject and thereby inducing the formation of thetissue in vivo.

As used herein “a composition composed of a synthetic polymer attachedto a naturally occurring protein or a portion thereof” refers to thepolymer-protein precursor molecule of the present invention which isdescribed hereinabove.

According to preferred embodiments of the present invention, the methodaccording to this aspect further comprises a step of cross-linkingfollowing administering the composition. Cross-linking can be performedas described hereinabove using non-toxic, non-hazardous agents and/orconditions.

According to preferred embodiments of this aspect of the presentinvention, the method further comprises administering to the subject amolecule capable of cross-linking the composition.

The phrase “molecule capable of cross-linking the composition” refers tothe cross-linking agent described hereinabove (e.g., PEG-DA).

It will be appreciated that the scaffold of the present invention can bealso used for ex vivo formation of a tissue. For example, thePEG-fibrinogen scaffolds of the present invention were shown to becapable of ex vivo nerve regeneration as illustrated in FIGS. 26-30 inExample 6.

Thus according to another aspect of the present invention there isprovided a method of inducing ex-vivo formation of a tissue.

The phrase “tissue” refers to part of an organism consisting of anaggregate of cells having a similar structure and function. Examplesinclude, but are not limited to, brain tissue, retina, skin tissue,hepatic tissue, pancreatic tissue, bone, nerve, cartilage, connectivetissue, blood tissue, muscle tissue, cardiac tissue brain tissue,vascular tissue, renal tissue, pulmonary tissue, gonadal tissue,hematopoietic tissue and fat tissue. Preferably, the phrase “tissue” asused herein also encompasses the phrase “organ” which refers to a fullydifferentiated structural and functional unit in an animal that isspecialized for some particular function. Non-limiting examples oforgans include head, brain, eye, leg, hand, heart, liver kidney, lung,pancreas, ovary, testis, and stomach.

As used herein, the phrase “ex vivo” refers to living cells which arederived from an organism and are growing (or cultured) outside of theliving organism, preferably, outside the body of a vertebrate, a mammal,or human being. For example, cells which are derived from a human beingsuch as human muscle cells or human aortic endothelial cells and arecultured outside of the body are referred to as cells which are culturedex vivo.

The method is effected by seeding the scaffold of the present inventionwith cells to thereby induce tissue formation.

The cells used by the present invention are capable of forming a tissue.Such cells can be for example, stem cells such as embryonic stem cells,bone marrow stem cells, cord blood cells, mesenchymal stem cells, adulttissue stem cells, or differentiated cells such as neural cells, retinacells, epidermal cells, hepatocytes, pancreatic cells, osseous cells,cartilaginous cells, elastic cells, fibrous cells, myocytes, myocardialcells, endothelial cells, smooth muscle cells, and hematopoietic cells.

The term “seeding” refers to encapsulating, entrapping, plating, placingand/or dropping cells into the scaffold of the present invention. Itwill be appreciated that the concentration of cells which are seeded onor within the scaffold of the present invention depends on the type ofcells used and the composition of scaffold used (i.e., molar ratiobetween the synthetic polymer and protein within the precursor moleculesand the percent of cross-linking molecule used).

It will be appreciated that seeding of the cells can be performedfollowing the formation of the hydrogel scaffold of the presentinvention (i.e., on the casted hydrogel scaffold) or on pre-castedhydrogels, i.e., by mixing the cells with the scaffold precursormolecules prior to cross-linking the scaffold. The concentration ofcells to be seeded on the hydrogels depends on the cell type and thescaffold properties and those of skills in the art are capable ofdetermining the concentration of cells used in each case.

It will be appreciated that following seeding the cells on the scaffold,the cells are preferably cultured in the presence of tissue culturemedium and growth factors.

For example, to induce the formation of a cartilage tissue, chondrocytesare seeded in the hydrogel scaffolds of the present invention (prior tocross-linking) at a concentration of approximately 15×10⁶ cell/ml,following which the seeded scaffold is placed in a casting frame at roomtemperature and is further subjected to photoinitiation as describedabove. Following hydrogel casting, the seeded scaffold is transferred toa Petri dish containing tissue culture medium supplemented with serum(e.g., 10% FBS) and/or 1% ITS (insulin, transferrin, and selenium,Sigma), and is incubated at 37° C. for a few weeks, during which theculture medium is replaced every other day.

Following seeding the cells in the scaffold of the present invention thescaffolds are routinely examined using an inverted microscope forevaluation of cell growth, spreading and tissue formation (see forexample FIGS. 18 a-e and 19 a-l).

Thus, the scaffold of the present invention which is formed in vitro, exvivo or in vivo can be used to induce tissue formation and/orregeneration and thus treat individuals suffering from tissue damage orloss.

Thus, according to another aspect of the present invention there isprovided a method of treating a subject having a disorder characterizedby tissue damage or loss.

As used herein the phrase “disorder characterized by tissue damage orloss” refers to any disorder, disease or condition exhibiting a tissuedamage (i.e., non-functioning tissue, cancerous or pre-cancerous tissue,broken tissue, fractured tissue, fibrotic tissue, or ischemic tissue) ora tissue loss (e.g., following a trauma, an infectious disease, agenetic disease, and the like) which require tissue regeneration.Examples for disorders or conditions requiring tissue regenerationinclude, but are not limited to, liver cirrhosis such as in hepatitis Cpatients (liver), Type-1 diabetes (pancreas), cystic fibrosis (lung,liver, pancreas), bone cancer (bone), burn and wound repair (skin), agerelated macular degeneration (retina), myocardial infarction, myocardialrepair, CNS lesions (myelin), articular cartilage defects(chondrocytes), bladder degeneration, intestinal degeneration, and thelike.

The phrase “treating” refers to inhibiting or arresting the developmentof a disease, disorder or condition and/or causing the reduction,remission, or regression of a disease, disorder or condition in anindividual suffering from, or diagnosed with, the disease, disorder orcondition. Those of skill in the art will be aware of variousmethodologies and assays which can be used to assess the development ofa disease, disorder or condition, and similarly, various methodologiesand assays which can be used to assess the reduction, remission orregression of a disease, disorder or condition.

The method is effected by implanting the scaffold of the presentinvention alone or following seeding such a scaffold with cells, or byadministering the scaffold units (i.e., the polymer-protein precursormolecules of the present invention) into the subject to thereby induceformation of the tissue and treat the disorder characterized by tissuedamage or loss.

It will be appreciated that the cells seeded on the scaffold for ex vivoformation of a tissue can be derived from the treated individual(autologous source) or from allogeneic sources such as embryonic stemcells which are not expected to induce an immunogenic reaction.

Following ex vivo tissue formation the seeded scaffold is implanted inthe subject. Those of skills in the art are capable of determining whenand how to implant the scaffold to thereby induce tissue regenerationand treat the disease. For example, if the disease to be treated isarticular cartilage the scaffold is seeded with chondrocytes andfollowing 14-21 days in culture the scaffold is preferably implanted inthe articular surface of the joint thereafter. Alternatively, thescaffold can be injected as a precursor solution, with or without cells,and polymerized directly in the site of the cartilage damage. Thepolymerized hydrogel fills the gap of the defect and initiates theregeneration of new cartilage tissue.

Compositions of the present invention may, if desired, be presented in apack or dispenser device, such as an FDA-approved kit, which may containone or more unit dosage forms (e.g., 100 mg) such as for personalizeduse containing the active ingredient (e.g., precursor molecules whichare not yet cross-linked such as PEGylated denatured fibrinogen) andoptionally sterile disposable means for delivery (e.g., syringe) and forillumination (e.g., illuminator covers). The pack may, for example,comprise metal or plastic foil, such as a blister pack. The pack ordispenser device may be accompanied by instructions for administration.The pack or dispenser device may also be accompanied by a notice in aform prescribed by a governmental agency regulating the manufacture,use, or sale of pharmaceuticals, which notice is reflective of approvalby the agency of the form of the compositions for human or veterinaryadministration. Such notice, for example, may include labeling approvedby the U.S. Food and Drug Administration for prescription drugs or of anapproved product insert. Compositions comprising a preparation of theinvention formulated in a pharmaceutically acceptable carrier may alsobe prepared, placed in an appropriate container, and labeled fortreatment of an indicated condition, as further detailed above.

It is expected that during the life of this patent many relevantpolymer-protein scaffolds will be developed and the scope of the termscaffold is intended to include all such new technologies a priori.

Additional objects, advantages, and novel features of the presentinvention will become apparent to one ordinarily skilled in the art uponexamination of the following examples, which are not intended to belimiting. Additionally, each of the various embodiments and aspects ofthe present invention as delineated hereinabove and as claimed in theclaims section below finds experimental support in the followingexamples.

EXAMPLES

Reference is now made to the following examples, which together with theabove descriptions, illustrate the invention in a non limiting fashion.

Generally, the nomenclature used herein and the laboratory proceduresutilized in the present invention include molecular, biochemical,microbiological and recombinant DNA techniques. Such techniques arethoroughly explained in the literature. See, for example, “MolecularCloning: A laboratory Manual” Sambrook et al., (1989); “CurrentProtocols in Molecular Biology” Volumes I-III Ausubel, R. M., Ed.(1994); Ausubel et al., “Current Protocols in Molecular Biology”, JohnWiley and Sons, Baltimore, Md. (1989); Perbal, “A Practical Guide toMolecular Cloning”, John Wiley & Sons, New York (1988); Watson et al.,“Recombinant DNA”, Scientific American Books, New York; Birren et al.(Eds.) “Genome Analysis: A Laboratory Manual Series”, Vols. 14, ColdSpring Harbor Laboratory Press, New York (1998); methodologies as setforth in U.S. Pat. Nos. 4,666,828; 4,683,202; 4,801,531; 5,192,659 and5,272,057; “Cell Biology: A Laboratory Handbook”, Volumes I-III Cellis,J. E., Ed. (1994); “Culture of Animal Cells—A Manual of Basic Technique”by Freshney, Wiley-Liss, N.Y. (1994), Third Edition; “Current Protocolsin Immunology” Volumes I-III Coligan J. E., Ed. (1994); Stites et al.(Eds.), “Basic and Clinical Immunology” (8th Edition), Appleton & Lange,Norwalk, Conn. (1994); Mishell and Shiigi (Eds.), “Selected Methods inCellular Immunology”, W. H. Freeman and Co., New York (1980); availableimmunoassays are extensively described in the patent and scientificliterature, see, for example, U.S. Pat. Nos. 3,791,932; 3,839,153;3,850,752; 3,850,578; 3,853,987; 3,867,517; 3,879,262; 3,901,654;3,935,074; 3,984,533; 3,996,345; 4,034,074; 4,098,876; 4,879,219;5,011,771 and 5,281,521; “Oligonucleotide Synthesis” Gait, M. J., Ed.(1984); “Nucleic Acid Hybridization” Hames, B. D., and Higgins S. J.,Eds. (1985); “Transcription and Translation” Hames, B. D., and HigginsS. J., Eds. (1984); “Animal Cell Culture” Freshney, R. I., Ed. (1986);“Immobilized Cells and Enzymes” IRL Press, (1986); “A Practical Guide toMolecular Cloning” Perbal, B., (1984) and “Methods in Enzymology” Vol.1-317, Academic Press; “PCR Protocols: A Guide To Methods AndApplications”, Academic Press, San Diego, Calif. (1990); Marshak et al.,“Strategies for Protein Purification and Characterization—A LaboratoryCourse Manual” CSHL Press (1996); all of which are incorporated byreference as if fully set forth herein. Other general references areprovided throughout this document. The procedures therein are believedto be well known in the art and are provided for the convenience of thereader. All the information contained therein is incorporated herein byreference.

Example 1 Generation of PEG-Fibrinogen Hydrogels

Tissue engineering scaffolds with controllable mechanical properties andadequate biofunctional signals were generated from PEG and fibrinogen.Briefly, denatured fibrinogen fragments were PEGylated withPEG-diacrylates, mixed with photoinitiator and exposed to UV light toform a hydrogel material in the presence of a cell suspension. Thedegradability of the PEG-fibrinogen scaffold was further tested byenzyme-mediated proteolysis, as follows.

Materials and Experimental Methods

Synthesis of PEG Diacrylate (PEG-DA)—PEG-diacrylate (PEG-DA) wasprepared from linear PEG, MW=4-kDa, 6-kDa, and 20-kDa (Fluka, Buchs,Switzerland), essentially as described elsewhere (Lutolf and Hubbell,2003; Elbert D L., et al., 2001). Briefly, acrylation of PEG-OH wascarried out under Argon by reacting a dichloromethane (DCM) (Aldrich,Sleeze, Germany) solution of the PEG-OH with acryloyl chloride (Merck,Darmstadt, Germany) and triethylamine (Fluka) at a molar ratio of 1-OHto 1.5-acryloyl chloride to 1.5-triethylamine (0.2 g PEG/ml DCM). Thefinal product is precipitated in the presence of ice-cold diethyl etherand dried under vacuum overnight. The degree of end-group conversion wasconfirmed by ¹H NMR and was found to be 97-99% (data not shown).

Cyanogen bromide cleavage of fibrinogen—Whole fibrinogen [Sigma-Aldrich,Steinheim, Germany, Cat # F8630, GenBank Accession No. AAC67562.1.(α-chain; SEQ ID NO: 1); GenBank Accession No. CAA23444.1 (β-chain; SEQID NO:2), and GenBank Accession No. CAA33562.1 (γ-chain; SEQ ID NO:3)]was dissolved in a solution of 70% formic acid containing 17 mg/mlCyanogen Bromide (CNBr) (Aldrich, Cat. # C9, 149-2) and incubatedovernight in the dark at 25° C. The cleaved fibrinogen fragments weredialyzed for 2 days at 4° C. in 50 mM phosphate buffered saline (PBS) atpH 7.4 with a twice-daily change of buffer to remove all the CNBr andformic acid from the solution. The dialyzed fragments were stored in PBSat 4° C. until they were subjected to PEGylation. Since grafting of PEGhas a detectable effect on the mobility of the protein in acrylamidegels (Kurfurst M M, 1992; Pomroy N C and Deber C M, 1998), visualizationof the PEGylated fibrinogen fragments was performed by loading thesamples on an SDS-PAGE followed by a Coomassie®-blue staining.

PEGylation of fibrinogen—To PEGylate the fibrinogen protein, tris(2-carboxyethyl) phosphine hydrochloride (TCEP-HCl) (Sigma) was added toa 7 mg/ml solution of fibrinogen in 50 mM PBS with 8 M urea (molar ratio68:1 TCEP to fibrinogen cysteins) and the solution was left shaking for15 min at 25° C. until fully dissolved. After dissolution of thefibrinogen, a solution of PEG-DA (250-300 mg/ml) in 50 mM PBS and 8 Murea was added to the fibrinogen and reacted overnight in the dark at25° C. The molar ratio of PEG to fibrinogen was 145:1 (linear PEG-DA, MW4-kDa, 6-kDa, and 20-kDa). The final PEGylated protein product wasprecipitated for 20 minutes at room temperature while stirring in 5×excess acetone (Frutarom, Haifa, Israel). The precipitated proteinsolution was centrifuged for 20 minutes at 5000 RPM (Sorvall GSA rotor)and the pellet was redissolved at 20 mg/ml protein concentration in PBScontaining 8 M urea. The PEGylated protein solution was then dialyzedfor 2 days at 4° C. against PBS containing 0.1% (v/v) glacial aceticacid (Frutarom) with twice-daily changes of PBS (Spectrum, 12-14-kDa MWcutoff). The dialyzed product was either used immediately or lyophilizedin a solution of 10% D-(+)-glucose (Riedel-deHaën, Germany) to improvesolubility upon redissolution. The lyophilized PEGylated product wasstored under Argon at −80° C. for up to six months.

Photo-polymerization of PEG hydrogels—The PEG-fibrinogen hydrogels weremade from a precursor solution of PEGylated fibrinogen (whole orcleaved). The precursor solution was made by solubilizing PEGylatedfibrinogen in 1 ml of 50 mM PBS, pH 7.4 and 25° C. to achieve a finalconcentration of 10, 15, or 20% polymer (w/v). The precursor solutionalso contained a PEG-DA cross-linking constituent at a molar ratio of1:2 PEG-DA to functional groups on the PEGylated fibrinogen. Theadditional PEG-DA was used to efficiently cross-link the PEGylatedprotein macromeres and minimize steric hindrances that result in poorgelation. The precursor solution was mixed with 10 μl of Igracure™2959photoinitiator solution (Ciba Specialty Chemicals, Tarrytown, N.Y.) in70% ethanol (100 mg/ml) and centrifuged for 5 min at 14,000 RPM. Thesolution was then placed into Teflon tubes (5-mm diameter and 20-mmlong) and polymerized under UV light (365 nm, 4-5 mW/cm²) for 15 minutesaccording to published protocols (Lum L Y et al., 2003). Followingpolymerization, the hydrogels were cut into 5-mm long sections formechanical testing. Control hydrogels were made by solubilizing 10, 15,or 20% (w/v) PEG-DA in 1-ml of 50 mM PBS with Igracure™ and thenpolymerizing under UV light as described above.

Mechanical properties testing—The compressive mechanical properties ofthe acellular PEG-fibrinogen hydrogels were evaluated using an Instron™5544 single column material testing system with Merlin software. Thestress-strain characteristics of 5-mm diameter plugs (5-mm long) weremeasured by constant straining (0.025 mm/sec) between two rigid,non-porous fixtures (unconfined). The material was strained to 30%strain and the force-displacement was recorded. The Merlin softwareautomatically converted the raw data into a stress-strain relationshipdescribing the material. The elastic modulus was determined directlyfrom the stress-stain data as the average slope of the lower portion ofthe stress-strain curve (<15% strain).

Biodegradation assay—To assess the rate of enzymatic hydrogeldegradation, Coomassie® brilliant blue G-250 dye (Aldrich) was bound tothe PEG-fibrinogen hydrogels and the release of the Coomassie® dye wasmeasured spectrophotometrically. Hydrogels were stained in 0.1%Coomassie® (w/v) overnight with gentle agitation, and destained for 1hour in destaining buffer. Since the Coomassie® dye binds to proteinswith very high affinity, following the initial staining, the hydrogelswere destained to release all unbound dye. The gels were thentransferred to multiwell plates and incubated with known concentrationsof Collagenase (Sigma) or trypsin (BD Biosciences, Sparks, Md.) in PBS.The enzymatic degradation by Collagenase and trypsin was correlated tothe release of Coomassie®-bound fibrinogen from the hydrogel networkfollowing a 24 hours incubation period. Samples of supernatant (350 μL)were transferred to a quartz cuvette and were measured using the UV/Visspectrophotometer (Hitachi Instruments Inc., USA). All data wasnormalized with the spectrophotometric measurements of hydrogelsdegraded completely by their respective enzyme.

Statistical Analysis—Statistical analysis was performed using thestatistical analysis features of Microsoft Excel. Data from at least twoindependent experiments were quantified and analyzed for each variable.Comparisons between multiple treatments were made with analysis ofvariance (ANOVA) while comparisons between treatments were made using atwo-tail student t-test with P<0.05 considered statisticallysignificant.

Experimental Results

PEGylation of fibrinogen can be performed using 4-kDa, 6 kDa and/or20-kDa PEG-DA—The PEG-fibrinogen hydrogels were prepared by covalentlybinding protein fragments to PEG-DA and cross-linking by UVphotoinitiation. FIG. 1 a illustrates the potential of the fibrinogenconstituent to be linked to fractionalized PEG. A Michael-type additionreaction (Lutolf M et al, 2001, Bioconjugate chem. 12(6):1051-6) wasused to form an ester bond between the free thiol groups in thefibrinogen cysteines and the acrylate end-groups on the PEG-DA (i.e.,PEGylation reaction) using varying molecular weight fragments of 4-kDa,6-kDa, and 20-kDa of PEG-DA.

To further confirm the presence of PEGylated fibrinogen fragments and todetermine the optimal conditions for the PEGylation reaction, thereaction products were subjected to an SDS-PAGE followed byCoomassie®-blue staining. As is shown in FIG. 2 a, following incubationof whole fibrinogen with a 4-kDa PEG-DA fragment, a time-dependent shiftto higher molecular weight fragments was observed. Thus, while followingone hour of incubation with the 4-kDa PEG-DA, many of the fragments wereelevated above 60 kDa, following two hours and/or overnight incubationsmost of the fragments were larger than 250 kDa. Moreover, as is shown inFIG. 2 b, when CNBr-cleaved fibrinogen was incubated with either the4-kDa or the 6-kDa PEG-DA fragments, a considerable shift to highermolecular weight fragments was observed within one hour of incubation.

These results clearly demonstrate that PEGylation of fibrinogen is moreefficient on CNBr-cleaved fibrinogen than on whole fibrinogen. Inaddition, these results show that PEGylation of fibrinogen is atime-dependent reaction.

The PEG-fibrinogen product exhibits high percentage of PEGylatedfibrinogen—Following an overnight PEGylation reaction the PEGylatedfibrinogen product is purified using acetone precipitation whichselectively precipitates the protein (i.e., PEGylated fibrinogen) fromthe excess of unreacted PEG-DA. It is worth mentioning that the additionof acetone at room temperature to the PEG-DA reaction [using an iodidesolution as described elsewhere (24)] did not result in theprecipitation of free, unreacted PEG-DA (data not shown). It will beappreciated that excess of unreacted PEG can theoretically becomeentangled with grafted PEG chains on the PEGylated protein and thusremain part of the PEGylated protein during precipitation. To confirmthe purity of the PEGylated fibrinogen, the dry weight of the totalPEGylated product was compared with the amount of total protein asmeasured using a Pierce BCA assay in a purified solution of PEGylatedfibrinogen. The dry weight of PEGylated protein should be the sum of theweights of the fibrinogen and the grafted PEG, assuming 100% PEGylation.The theoretical protein fractions for PEGylated fibrinogen using 4-kDa,6-kDa, and 20-kDa PEG (assuming 100% PEGylation and 100% purity) are59%, 49%, and 22%, respectively. The results indicate a protein fractionof 45±5.0% using the 4-kDa PEG, 39±3.4% using the 6-kDa PEG, and 36±1.5%using the 20-kDa PEG. The difference between the theoretical and themeasured protein factions can be attributed to several factors,including excess of unreacted PEG, partial PEGylation, and/or error inthe BCA measurements. It is worth mentioning that excess of unreactedPEG was not visible in SDS-PAGE stained with iodine-acetate (data notshown), thus confirming the observation that the fibrinogen is highlyPEGylated after an overnight reaction with PEG-DA. Thus, these resultssuggest that the deviation from the expected protein fractions resultfrom either partial PEGylation or errors in the determination of proteinconcentrations using the BCA assay.

Cyanogen bromide cleavage of fibrinogen facilitates the preparation ofPEGylated fibrinogen solutions—Since the combination of the highlyhydrophilic PEG to the partially hydrophobic fibrinogen can result in ahighly hydrophobic protein core, the fibrinogen molecule was denaturedprior to being subjected to the PEGylation reaction. Such denaturationminimizes the formation of the hydrophobic core after PEGylation andsignificantly improves the solubility of the product. Denaturation ofthe fibrinogen molecule was accomplished using CNBr, a proteolyticmolecule which chemically cleaves adjacent to methionine peptides in thefibrinogen sequence. CNBr cleavage of bovine fibrinogen (3 partsfibrinogen to 2 parts CNBr) results in 30 fragments (F1-F30) ranging insize from 35-kDa to 0.1-kDa each. Eight of these fragments contain twoor more unpaired thiol groups which are capable of contributing to thestructure of the hydrogel network (see Tables 1 and 2, hereinbelow).FIG. 2 a demonstrates the presence of small fragments of fibrinogenbefore and after PEGylation using SDS-PAGE. In addition, the PEGylatedCNBr-cleaved fibrinogen fragments exhibit the same PEGylation efficiencyand protein purity as the whole PEGylated protein after acetoneprecipitation (data not shown). Thus, using the method of the presentinvention, PEGylated fibrinogen solutions with concentrations of up to300 mg/ml were prepared under non-denaturing conditions.

TABLE 1 Fibrinogen fragments and cysteines from whole fibrinogenFragments Alpha beta gamma Total M.W. 65 53.3 47.6 165.9 Cysteines 8 1110 29

TABLE 2 Fibrinogen fragments and cysteines from multi thiol fragments ofcleaved fibrinogen Fragments F1 F2 F3 F4 F7 F8 F15 F16 Total M.W. 32.424.6 12.1 10 7 6.3 3.8 3.8 100 Cysteines 2 4 3 4 2 2 4 4 25

PEG-fibrinogen hydrogels are highly elastic—To confirm formation ofhydrogels from the PEGylated protein precursors, the compressivemechanical properties of the material were measured using the Instron™single column material testing system. These measurements confirmed theformation of acellular PEG-fibrinogen hydrogels (and PEG-PEG controls)with varying amounts of polymer (10%, 15%, and 20% w/v) and differentmolecular weights of PEG (4-kDa and 6-kDa, 20-kDa) as well as theformation of hydrogels which are made of either whole fibrinogen (whole)or CNBr-cleaved fibrinogen (cleaved). Further measurements of thestress-strain characteristics of the PEG hydrogels demonstrated that thestress-strain characteristic of both PEG-PEG and PEG-fibrinogenhydrogels is non-linear and highly dependent on the molecular weight ofthe PEG precursor (FIGS. 3 a and b). In addition, as is shown by thestress-strain graphs (FIGS. 3 a-b), the modulus of elasticity observedin the PEG-fibrinogen hydrogels (in the range of 0.12-0.14) wassignificantly lower than that of the PEG-PEG hydrogels (in the range of2.5-3.21). In addition, the elastic modulus (as determined from thestress-strain curve) was found to be dependent on the percent polymer,the molecular weight of the PEG precursor, and the fibrinogen backbone(FIGS. 4 a-c). In general, the elastic modulus of the PEG-PEG hydrogels(at any given percent polymer concentration) was found to besignificantly higher than that of the PEG-fibrinogen hydrogels (n=5,p<0.05). Likewise, the elastic modulus of PEG-fibrinogen hydrogels madewith whole fibrinogen was found to be significantly different than thatof the hydrogels made with cleaved fibrinogen (n=5, p<0.05).

Biodegradation of PEG-fibrinogen hydrogels is dependent on the MW of thePEG component and the protein backbone—Hydrogel biodegradation wasquantified by subjecting colorimetrically-labeled pure PEG-fibrinogenhydrogels (15% w/v PEG-DA, MW 6-kDa) to varying concentrations ofproteases (e.g., Collagenase or trypsin) and assessing the dissolutionof the gels. Thus, as the hydrogel dissolves, fragments released into anoverlaying buffer are quantified using a spectrophotometer. As is shownin FIG. 5 a, degradation of the PEG-fibrinogen hydrogels was affected byboth the MW of the grafted PEG and the molecular structure of theprotein backbone of the hydrogel. Thus, changing the PEG MW from 6-kDato 20-kDa resulted in accelerated degradation in the presence of 0.05mg/ml trypsin (n=6, p<0.05) but not in the presence of 0.5 mg/mlCollagenase. Likewise, the degradation of the cleaved fibrinogenhydrogels in 0.5 mg/ml Collagenase was significantly higher than that ofthe whole fibrinogen hydrogels (n=6, p<0.05). As is shown in FIG. 5 b,increasing concentrations of either trypsin or Collagenase resulted inhigher percent degradation of the hydrogels. Similar results wereobtained by analyzing the degraded PEGylated fibrinogen hydrogels onSDS-PAGE (data not shown).

Altogether, these results demonstrate the generation of uniquePEG-fibrinogen hydrogels which provide distinct advantages over otherscaffold materials: (i) the mechanical properties of the PEG-hydrogelsare highly malleable; (ii) the biological functionality is maintained bythe protein backbone of the polymeric network; and (iii) the elasticmodulus of the PEG-fibrinogen hydrogel is dependent on the molecularweight of the PEG constituent and proportional to the percent polymericcomposition.

Analysis and Discussion

The use of protein-based synthetic materials is a novel approach todesigning the “next-generation” of hydrogel scaffolds for tissueengineering (16, 23, 26-30). These materials are capable of promotingcell growth and exhibit proteolytic degradability via their biologicaldomains while still providing exacting mechanical properties based ontheir synthetic composition. A number of these hybrid hydrogel materialsare currently in use as ingrowth matrices or cell culture substrateswhich take advantage of biologically active oligopeptides in thematerial backbone that mimic the properties of natural tissue (17, 18).While these materials satisfy the general criteria of biofunctionality,the small oligopeptides provide only a fraction of the bioactive signalspresent in the natural extracellular matrix (ECM). To overcome thislimitation, some approaches use recombinant DNA technologies to createan engineered protein-like backbone with inherent bioactivity (16),while other approaches add bioactive growth factors that are eithercovalently (23, 27, 31) or non-covalently (32) immobilized into thematerial.

The hydrogel material of the present invention contains a naturalprotein backbone onto which di-functional PEGs are covalently bound andcross-linked together using photo-polymerization. The protein backboneis comprised of alpha, beta, and gamma fragments of denaturedfibrinogen. These fibrinogen fragments are inherently bioactive withproteolytically sensitive sequences, cell adhesion motifs, and othercell-signaling sequences (14). Free thiol groups present in unpairedcysteine residues in the denatured fibrinogen fragment are covalentlyconjugated by a Michael-type addition reaction to an unsaturated doublebond on functionalized PEG-DA. The hydrogel material of the presentinvention uses denatured fibrinogen since it consists of a large numberof free thiol groups (i.e., unpaired cysteine residues) that can reactwith PEG-DA. After PEGylation, unreacted acrylates on the di-functionalPEGs are used to cross-link the fibrinogen backbone into a hydrogelnetwork using photo-polymerization.

The molecular structure of the hydrogels is highly influenced by thedegree of cross-linking. In addition to the percent polymer, the othercrucial determining factor of the hydrogel cross-linking is the ratio ofreactive acrylate end-groups on the PEG per MW of protein (acrylates perkDa of PEGylated protein). Assuming that each PEG is bound to fibrinogenwith one acrylate end-group, fully PEGylated fibrinogen contains 29acrylate groups for each molecule of protein, or approximately 0.1acrylate/kDa for 4-kDa PEGylated fibrinogen, 0.086 acrylate/kDa for6-kDa PEGylated fibrinogen, and 0.039 acrylate/kDa for 20-kDa PEGylatedfibrinogen. A pure PEG-DA solution contains 0.5 acrylates/kDa for 4-kDaPEG, 0.33 acrylates/kDa for 6-kDa PEG, and 0.1 acrylates/kDa for 20-kDaPEG. If the ratio of acrylates to MW is too low, the polymer solutionwill not form a continuous hydrogel. For this reason, the experimentsperformed in the present study utilized high concentrations of PEGylatedfibrinogen (≧10%).

The PEGylated fibrinogen requires substantial effort to solubilize atthese high concentrations. Since the solubility of the PEGylatedfibrinogen is highly affected by the grafting of the PEG chains;presumably because of the formation of a hydrophobic protein complex inthe presence of a highly hydrophilic PEG graft, the fibrinogen proteinused by the present invention utilized an irreversibly cleaved by CNBr.Thus, PEGylation of cleaved fibrinogen results in a highly solubleprotein precursor for hydrogel formation.

The mechanical properties of the cleaved and whole PEG-fibrinogenhydrogels was characterized using the stress-strain behavior underquasi-static uniaxial unconfined compression. The PEG-fibrinogenmaterial demonstrates the typical non-linear stress-straincharacteristics of a polymeric material, similar to the PEG-PEG controlgels. The composite behaves like a viscoelastic solid with minimalhysteresis under repetitive cyclic loading (data not shown). Thematerial stiffness, as determined from the modulus of elasticity, isdirectly proportional to the percent polymer composition. It will beappreciated that the material stiffness of polymeric hydrogels can bedirectly attributed to the degree of cross-linking. In thePEG-fibrinogen polymer, the amount of functional groups available forcross-linking is proportional to the polymer concentration. Thus, themodulus of the material is directly proportional to the amount polymerin the hydrogel.

The relationship between the elastic modulus of the hydrogel and themolecular chain length of the grafted PEG constituent is ambiguous.While pure PEG hydrogels exhibit a proportional relationship between thematerial modulus and the MW of the PEG, the addition of the protein intothe hydrogel network can have a profound impact on this relationship,depending on the relative size of the two polymers. In the case of wholePEGylated fibrinogen hydrogels, where the PEG is significantly smallerthan the protein, the impact of PEG MW on the modulus is morepronounced. In contrast, the relationship between PEG MW and the modulusof the hydrogels is less pronounced when the hydrogels are made ofcleaved protein. In either case, it is difficult to resolve thedifference between 4-kDa PEG and 6-kDa PEG when comparing the elasticmoduli.

The mechanical properties data demonstrate a significant difference inthe stiffness of hydrogels made from pure PEGylated fibrinogen versuspure PEG hydrogels. Irrespective of the PEG MW, the PEGylated fibrinogenhydrogels are always less stiff than the pure PEG hydrogels. This can beattributed to the fact that pure PEG hydrogels (made with similar weightpercent of polymer) contain nearly 5 times more functional groupsavailable for cross-linking than the PEGylated hydrogels. Thediscrepancy in cross-linking sites arises from the fact that thefibrinogen constituent accounts for more than half the weight of thepolymer but does not contain inherent cross-linking sites, and eachgrafted PEG on the PEGylated fibrinogen only has a single functionalgroup available for cross-linking, in contrast to two functional groupson free PEG-DA. Additional cross-linking limitations may arise becauseof steric hindrances caused by the bulky PEGylated fibrinogen molecules.

The addition of free PEG-DA to the PEGylated fibrinogen hydrogelsintroduces cross-linking sites to the hydrogel network and alters thematerial stiffness accordingly. Therefore, the balance between PEGylatedfibrinogen and the additional free PEG-DA can be used to modulate thestiffness of the hydrogels without compromising the biofunctionaldomains of the fibrinogen constituent. It is important to note that theaddition of free PEG to the PEGylated fibrinogen hydrogel can have asignificant impact on the degradability of the hydrogels in the presenceof cell-secreted proteases (data not shown).

Based on these findings, it is clear that the molecular chain length andpercent composition represent two independent parameters to control themechanical properties of the material. An additional parameter which canalter the mechanical properties of the material is the number ofcross-linking sites on each grafted PEG molecule. As eluded to earlier,the linear grafted PEG contains only one functional group forcross-linking the hydrogel; however, PEG molecules containing severalfunctional groups such as star-PEGs can also be grafted onto thefibrinogen to form hydrogels. While star-PEG was not used as part of thecurrent investigation, future studies aim to increase the stiffness ofthe PEG-fibrinogen hydrogels using 4-arm and 8-arm star-PEG precursors.

The fibrinogen backbone of the hydrogel material provides proteolyticsensitivity via naturally occurring substrates for fibrinolysis.Fibrinolysis is a physiological process whereby the native fibrinmolecule is proteolytically dismantled by serine proteases. In theory,the fibrinogen backbone is cleaved in the presence of activatedproteases, resulting in complete dissolution of the PEG-fibrinogenhydrogel. The results shown in the present study demonstrate that purePEG-fibrinogen hydrogels are proteolytically degradable while thePEG-PEG controls are not susceptible to proteolysis (data not shown).

The degradation data reveals several other interesting patternsregarding the proteolytic degradation of cleaved and whole fibrinogenhydrogels in the presence of trypsin or Collagenase after 30 minutes.Cleaved fibrinogen presents less substrate for degradation therebyincreasing the normalized dissolution of the hydrogel in Collagenase(n=5, p<0.05). In trypsin, the cleaved and whole fibrinogen hydrogelsdegrade almost identically. This is likely explained by the observationthat 0.05 mg/ml trypsin is saturated with substrate whereas 0.5 mg/mlcollagenase is not saturated with substrate. Evidently, the MW of thePEG constituent also affects the degradation results. In trypsin, the20-kDa PEG hydrogels are significantly more degraded than 6-kDa PEGhydrogels after 30 min, while in Collagenase there is no significantdifference in the degradation between the two conditions (n=6, p<0.05).The 20-kDa PEG hydrogels are comprised of less fibrinogen which resultsin faster hydrogel dissolution. This can also be explained by theobservation that 0.05 mg/ml trypsin is saturated with enzyme and cannotdegrade the 20-kDa PEGylated fibrinogen hydrogels faster. Future studieswill examine more aspects of proteolytic degradability of thePEG-fibrinogen material, including degradation kinetics and fibrinolysisin the presence of plasmin.

Example 2 The PEG-Fibrinogen Hydrogels Support Cell Spreading andExtension

To test the capacity of the PEG-fibrinogen hydrogels to promote theexpansion and differentiation of cell cultures, cell of bovine aorticendothelial cells and smooth muscle cells were cultured on variousPEG-based hydrogels, as follows.

Materials and Experimental Methods

In vitro cell-culture studies—Bovine aortic smooth muscle cells (BSMCs)from young donors were isolated and cultured according to a modifiedprotocol of Oakes et al., 1982. The BSMCs were cultured up to 6^(th)passage in Dulbecco's Modified Eagle Medium (DMEM) (Gibco, U.K.)containing 10% fetal bovine serum (FBS) (Biological Industries, Israel),1% penicillin-streptomycin (Biological Industries), and 1% L-glutamine(Gibco). PEG hydrogels containing BSMCs were made by mixing a PBS cellsuspension and PEGylated fibrinogen precursor solution containingIgracure™ photoinitiator to make a 10% (w/v) solution with 1.5×10⁶cells/ml. Aliquots of 100 μl of the suspension were added into wells ina flat-bottom 96-well plate and placed under UV-light (4-5 mW/km²) for 5min in a laminar flow hood. DMEM culture medium (containing 10% FBS) wasadded immediately to the polymerized hydrogels and changed daily(100-μl/well).

Viability studies of cells following exposure to UV light—To verify thatthe exposure to UV light did not damage the cells in the hydrogel,bovine aortic endothelial cells (BAECs) were isolated and cultured up to6^(th) passage according to published protocols (Remuzzi A., et al.,1984). The BAECs were seeded and cultured on top of 1-mm thick PEGhydrogels in 24-well plates as described elsewhere (Seliktar D., et al.,2004). The seeding density of BAECs was 30,000 cells/cm². BAECs andBSMCs were monitored daily using a Nikon TE2000 phase-contrastmicroscope and digitally imaged with a digital CCD camera (Jenoptik,Germany).

Histological evaluations—The cellularized PEG-fibrinogen hydrogels werefrozen in liquid nitrogen-cooled 2-methylbutane (J. T. Baker,Phillipsburg, N.J.) and cut in 7 μm-thick sections using a cryostat. Thesections were fixed onto glass microscope slides with ice-cold acetoneand stained with Hematoxylin and Eosin (H&E) to visualize cellmorphology. Cell morphology was documented using a Nikon lightmicroscope (TS-100) connected to a digital imaging workstation (SonyCorporation, Japan).

Experimental Results

PEG-fibrinogen hydrogels support cell spreading and attachment—To testthe capacity of the PEG-fibrinogen hydrogels to supportthree-dimensional spreading and attachment, endothelial (BAECs) orsmooth muscle (BSMCs) cells were cultured on the surface of or insidethe PEG hydrogels. BAECs were seeded on the surface of the hydrogels ata concentration of 30,000 cells/cm² and 24 hours following cell seedingthe degree of cell attachment and spreading was evaluated using phasecontrast microscopy. As is shown in FIGS. 6 a and d, while BAECs grownon PEG-fibrinogen hydrogels exhibited significant cell attachment andspreading (FIG. 6 a), BAECs cells grown on PEG-PEG hydrogels were roundand devoid of any visible spreading (FIG. 6 d). These resultsdemonstrate the superior capacity of the PEG-fibrinogen to support celladhesion or spreading over the PEG-PEG hydrogels. The capacity of thePEG-fibrinogen hydrogels to support three-dimensional spreading andattachment of cells was further tested dispersing BSMCs cells into theprecursor solution (1.5×10⁶ cells/ml, 4-kDa PEG, 10% polymer) prior tophoto-polymerization. After assembly, the hydrogel network containedhomogeneously distributed BSMCs with round morphology. However,following 24 hours of culturing, BSMCs cultured inside thePEG-fibrinogen hydrogels formed stable adhesions, processes, andcellular extensions (FIGS. 6 b-c). In addition, PEGylated fibrinogenhydrogels made with cleaved fibrinogen also supported the adhesion andextension of BSMCs inside PEGylated hydrogels (data not shown). Incontrast, BSMCs cultured inside PEG-PEG hydrogels were round and devoidof any visible cell extensions (FIG. 6 e). Further histologicalevaluation of the cultured BSMCs within the PEG hydrogels confirmed theobservation that BSMCs are extended within the PEG-fibrinogen hydrogelnetwork (FIG. 7 a), but not within the PEG-PEG control hydrogels (FIG. 7b).

Analysis and Discussion

The PEG hydrogel scaffold without the fibrinogen backbone is completelydevoid of biofunctional domains for cell culture. The fibrinogenbackbone provides at least two biofunctional characteristics to thehydrogel material: proteolytic sensitivity and cell adhesivity. Withregards to the latter, the PEG-fibrinogen hydrogels support theattachment and spreading of endothelial cells in the presence of serumproteins whereas PEG-PEG controls are not able to support cellattachment. Endothelial cell-surface adhesion molecules can thereforeattach either directly to adhesion domains on the fibrinogen backbone orto other serum proteins that interact non-specifically with thefibrinogen. Once attached, cells are capable of proteolyticallytunneling through the hydrogel network with the help of cell-secretedenzymes such as Collagenase. This is best exemplified with smooth musclecells that are three-dimensionally entrapped inside the hydrogelmaterial after photopolymerization and begin to form clusters of cellsafter 24 hours in culture. Phase contrast micrographs and histologicalcross-sections confirm that BSMCs form flagella-like extensions whichenable their migration inside the hydrogel network. In this regards,there are no observed differences between the whole and cleavedfibrinogen hydrogel scaffolds. In contrast, BSMCs remained round andhomogeneously dispersed and cell extensions are not observed in thenon-degradable PEG-PEG controls.

Thus, the biological domains in the fibrinogen backbone provideattachment motifs for endothelial cell and smooth muscle cell adhesionas well as proteolytic sensitivity for biodegradation. Smooth musclecells demonstrate the ability to proteolytically penetrate through thehydrogel material and form interconnecting networks of cells. Thus, thescaffolds of the present invention are novel, biodegradable and highlysuitable for cultivating cells in a 3-D environment for tissueregeneration therapies.

Example 3 In Vivo Regeneration of Bone Using the PEG-Fibrinogen Gelrin™Scaffold

To test the potential of the PEG-fibrinogen scaffold material tofacilitate tissue regeneration, a critical size tibial defect wasintroduced in rats and the Gelrin™ scaffold was implanted at the site ofsurgery (tibia diaphysis).

Materials and Experimental Methods

Animals—Female Sprague-Dawley rats (age 3-4 months) were adapted toanimal cage life for 5 days prior to the experiment. The weight of theanimal was monitored during this period to ensure stability and properadaptation. The animals were fed and watered daily without restrictions.

Introduction of a critical size tibia defect—The animals wereanesthetized with a combination of Ketamine (120 mg/kg) and Xylazine (17mg/kg). During the surgical procedure the animals were placed on a warmplate to maintain body temperature (and prevent hypothermia). The righttibia was shaved and wiped with polydine tincture solution. Themid-portion of the right tibia was exposed from the anterior medial sideby longitudinal incision (FIG. 8 a). An external fixator was placedproximal and distal to the mid-section of the tibia (FIG. 8 b). Twoneedles were drilled into the proximal tibia (21G) and distal tibia(23G) and connected to two external fixators (screws) to form a stablefixation of the bone. A 10-mm gap was excised using a disk saw in theportion between the proximal and distal needles of the fixators (FIG. 8c). The fibula was not osteotomized.

Implantation of a PEG-fibrinogen (Gelrin™) scaffold—A PEG-fibrinogenplug (5-mm diameter and 10-mm long) was inserted into the defect siteand the surrounding connective tissue was wrapped around to secure theplug into place (FIG. 8 d). The incisional wound was sutured using nylonsurgical thread. The animal was given prophylactic antibiotics(ampicilline 0.1 gram/100 gram). Immediately following surgery, theanimal was x-rayed and further evaluated weekly by x-ray screening. Theanimal was free to move about the cage during the entire post-operativefollow-up period. At the end of the 2-month evaluation period, theanimal was sacrificed with CO₂ and the right tibia was harvested forhistology and mechanical testing.

Experimental Results

Introduction of a critical size rat tibia defect is known to result inup to 20% mortality. A critical rat tibia defect was introduced into 25rats (FIGS. 8 a-d), of which 17 were further subjected to scaffoldimplantation using the PEG-fibrinogen or PEG-PEG hydrogels (see Table 3,hereinbelow, for representative rats). Four to seven weeks followingsubjecting rats to critical size tibia defects the rats were sacrificedand the tibias were evaluated for the presence, extent and location oftissue ingrowth and new bone formation. As is shown in FIGS. 9 a-b, fiveweeks following surgery a new bone can be seen in the Gelrin™-implantedrat tibia (FIG. 9 b) but not in control rats (FIG. 9 a).

TABLE 3 Scaffold implantation in critical size rat tibia defect Days RatSurgery post- # Date Leg Hydrogel Composition operation 5 24/5/04 LeftPEG-fibrinogen (cleaved, 4-kDa) 42 15% W/V total; 10 % PEG-Fibrinogenand 5% PEG-DA 6 24/5/04 Right PEG-fibrinogen (cleaved, 4-kDa) 35 15% W/Vtotal; 10 % PEG-Fibrinogen and 5% PEG-DA 7 16/8/04 Right PEG-fibrinogen(whole, 10-kDa) 49 1.75% PEG-fibrinogen (Gelrin ™) and 3% PEG-DA 816/8/04 Right PEG-fibrinogen (whole, 10-kDa) 1.75% PEG-fibrinogen(Gelrin ™) and 3% PEG-DA 9 16/8/04 Right PEG-fibrinogen (whole, 10-kDa)35 1.75% PEG-fibrinogen (Gelrin ™) and 3% PEG-DA 11 27/7/04 RightPEG-only (4-kDa) 10% W/V 41 12 27/7/04 Right PEG-only (4-kDa) 10% W/V 4113 27/7/04 Right PEG-only (6-kDa) 15% W/V 41 14 27/7/04 Right PEG-only(6-kDa) 15% W/V 41 15 27/7/04 Right Empty - control 41 16 27/7/04 RightEmpty - control 41 21 13/7/04 Right PEG-fibrinogen (whole, 10-kDa) 1.75%PEG-fibrinogen (Gelrin ™) and 3% PEG-DA 22 13/7/04 Right PEG-fibrinogen(whole, 10-kDa) 1.75% PEG-fibrinogen (Gelrin ™) and 3% PEG-DA 23 27/9/04Right PEG-fibrinogen (whole, 10-kDa) 1.75% PEG-fibrinogen (Gelrin ™) and3% PEG-DA 24 27/9/04 Right PEG-fibrinogen (whole, 10-kDa) 1.75%PEG-fibrinogen (Gelrin ™) and 3% PEG-DA Table 3: Representative in vivoexperiments performed following the introduction of critical size rattibia defect. Rats were implanted with either PEG—PEG (PEG-only) orPEG-fibrinogen hydrogels. The PEG-fibrinogen hydrogels were preparedfrom PEG-fibrinogen precursors and PEG-DA cross-linkers. ThePEG-fibrinogen precursors were prepared from either whole fibrinogen(Gelrin ™) or from CNBr-cleaved fibrinogen.

FIGS. 10-17 demonstrate the presence of normal bone and cartilage tissuefive weeks following introduction of a critical size tibia defect andGelrin™-implantation. Thus, tibias of Gelrin™-treated rats exhibitedwell-vascularized, orientated and densely textured fibrous tissue (FIG.13), osteonal healing (FIG. 11), and Haversian systems with smallcentral canals containing blood vessels (FIG. 10).

Example 4 Biodegradability of PEG-Fibrinogen Hydrogels is Controllablevia the Synthetic PEG Constituent

Biodegradability of PEG-fibrinogen hydrogels is reduced by the additionof free PEG-DA to the hydrogels—To further characterize the effect ofthe concentration of the cross-linking molecule of the present invention[i.e., PEG-DA (free PEG)] on the hydrogel biodegradability, 1×10⁶ smoothmuscle cells per ml were seeded inside Gelrin™ hydrogels (1.75%PEG-fibrinogen, 10-kDa PEG) containing various concentrations of PEG-DA.The seeded hydrogels were cultured for 48 hours, following which theability of the cells to attach and spread inside the hydrogels wasevaluated using an inverted microscope. As is shown in FIGS. 18 a-e,while cells cultured in the pure Gelrin™ matrix (in the absence of freePEG-DA) exhibited cell extensions and significant spreading, cellscultured in hydrogels fabricated with increasing concentrations of freePEG (i.e., 0.5-2%) were devoid of cell extensions and were more rounded.These results demonstrate that excess of a cross-linking molecule(PEG-DA) in the Gelrin™ matrix reduces the cell-mediated degradabilityand cell-penetration through the hydrogel.

Biodegradability and cell extension is decreased in Gelrin™ hydrogelsconsisting of more than 1% of a cross-linking molecule (PEG-DA)—Tofurther test the effect of the cross-linking molecule (i.e.,functionalized PEG-DA) on cell spreading through the PEG-fibrinogenscaffold of the present invention, Gelrin™ matrices (1.75%PEG-fibrinogen, 10-kDa PEG) were prepared as pure hydrogels (in theabsence of free PEG-DA cross-linker) or in the presence of 1 or 2% offree PEG-DA and cell clusters composed of highly compacted smooth musclecells inside a collagen gel matrix (5×10⁶ cells in one mg collagencluster). The clusters were seeded in the center of each hydrogel byplacing the cell mass into the Gelrin™ matrix prior to polymerizationand forming the scaffold around the tissue so that it encapsulates thetissue mass from all sides. The degree of cell extension was monitoredfollowing 1, 2, 4, and 7 days in culture using phase microscopy. As isshown in FIGS. 19 a-l, significant cell extensions were observed in pureGelrin™ hydrogels beginning following one day in culture. In contrast,cell extensions were relatively decreased in smooth muscle cell clusterscultured inside Gelrin™ hydrogels which were fabricated in the presenceof 2% free PEG-DA.

Altogether, these results demonstrate that the biodegradability and cellspreading through the scaffolds depends on the degree of the hydrogelcross-linking; higher concentrations of a cross-linking molecule(PEG-DA) inversely correlates with biodegradability of the scaffoldhydrogel of the present invention.

Example 5 In Vivo Osteogenesis Mediated Bypegylated FibrinogenDegradation Products

The osteoinductive properties of denatured PEGylated fibrinogendegradation products in osseous regeneration were studied in asite-specific bone defect. In order to identify the optimal compositionof PEG and fibrinogen required for synchronized hydrogel degradationduring the defect healing response in the osseous environment, PEGylatedfibrinogen hydrogels were prepared with three compositions of syntheticPEG constituent, each providing the material with slightly differentsusceptibility to proteolytic degradation.

Materials and Experimental Methods

Introduction of a critical size tibia defect—Tibia defects wereintroduced into 24 female Sprague-Dawley rats as described in Example 3,except that a 7 mm gap and not a 10 mm gap was excised from the tibia.

Implant Fabrication—Acellular cylindrical plugs were cast in 3-mmdiameter silicon tubes using 88 μl aliquots of PEG-fibrinogen precursorby a radical chain polymerization reaction of acrylate end groups.Additional PEG-DA (3% or 5% w/v) was added to the precursor solution inorder to increase the number of cross-links and to reduce thesusceptibility of the protein backbone to proteolytic degradation. Thefinal ratio of PEG to fibrinogen monomer was roughly 25:1, 100:1, and150:1 for the 0% PEG-DA, 3% PEG-DA, and 5% PEG-DA, respectively. Theprecursor solution was mixed with 0.1% (v/v) photoinitiator stocksolution made of 10% w/v Irgacure™2959 (generously donated by CibaSpecialty Chemicals, Tarrytown, N.Y.) in 70% ethanol and deionizedwater. The solution was placed under a UV light (365 nm, 4-5 mW/cm²) for5 minutes to polymerize. The pre-cast hydrogels were stored in 50 mM PBScontaining 2% penicillin-streptomycin (Biological Industries, Israel)for 5 hrs prior to implantation.

Implantation of PEG-fibrinogen plugs—PEG-fibrinogen plugs (3-mm diameterand 7-mm long) were inserted into the defect sites as described inExample 3 hereinabove and radiographed shortly after the surgery andthereafter evaluated at weekly intervals by x-ray screening. Altogether,6 rats were not treated i.e. boney gaps were left empty, 6 rats weretreated with PEG fibrinogen plugs comprising no additional PEG-DA(treatment 1), 6 rats were treated with PEG fibrinogen plugs comprising3% additional PEG-DA (treatment 2) and 6 rats were treated with PEGfibrinogen plugs comprising 5% additional PEG-DA (treatment 3). Asummary of the characteristics of the three treatment types is providedin Table 4 hereinbelow.

TABLE 4 Summary of Treatment Cohorts Degradation Degradation Rate inRate in Composition Trypsin Collagenase Degra- PEG:Fibrinogen (% Wt (%Wt Group dation ratio Loss/min^(1/2)) Loss/min^(1/2)) Control N/A Emptygap N/A N/A 1 Fast 25:1 8.043 9.730 2 Inter- 75:1 1.186 2.045 mediate 3Slow 150:1  0.872 1.007

Histological Analysis: Following a final radiographic evaluation, theright tibia of each rat was carefully excised in its entirety. Thesamples were fixed in buffered, neutral 10% formalin solution for 10days and then decalcified in 10% formic acid for 10 days. The specimenswere trimmed so as to include the implant site and the adjacent bonetissue on either side of the defect. Following rinsing in PBS, thespecimens were dehydrated in increasing concentrations of ethanol indeionized water (70% to 100%). They were embedded in extra-largeparaffin blocks. The latter were sectioned at 6 μm, fixed onpoly-1-lysine coated glass slides, and stained with hematoxylin andeosin (H&E).

Experimental Results

Osteogenesis—Newly formed bone in the site-specific defects of thetibiae was radiographically observed as early as three weekspostoperatively. When compared to control rats, large amounts of newbone were apparent in the defects of the treatment-2 animals (3%additional PEG-DA) by 5 weeks (FIG. 20). This contrasted with the lackof radiographically detectable bone in the defects of the treatment-1(0% additional PEG-DA), treatment-3 (5% additional PEG-DA), and controlrats. The histological examination confirmed that the rats treated withthe intermediate-degrading hydrogel (treatment-2) exhibited the mostextensive and widespread osteoneogenesis in and nearby the defect site.The longitudinally, H&E-stained sections of the tibiae revealed that theextent of regenerated bone in the site-specific defects ranged frompartial to total bridging of the gap (FIGS. 21 a-c). Osteoneogenesis wasobserved at both the endosteal and subperiosteal aspects. When observedunder polarized light, the birefringent pattern of the preexistinglamellar-fibered cortical bone sharply contrasted to that of thewoven-fibered boney trabeculae, characteristic of newly depositedosseous tissue. The newly formed subperiosteal bone at the osteotomysites was contiguous with the boney trabeculae, which were for the mostpart rimmed by cuboidal osteoblasts on their inner front. Theendosteally formed bone was as well continuous with newly formedtrabeculae, which extended into the defect site. Randomly scatteredadipocytic islands were present in between the trabeculae of the newlyformed woven-fibered bone. In but a few samples, the medullary cavitycontained some fibrous tissue proximal to the osteotomy site.

In those cases in which there was total osseous bridging of thesite-specific defect, the implant had been entirely resorbed andreplaced by lamellar-fibered bone with an atypical pattern of Haversian,i.e., osteonal bone. The newly formed bone was uninterrupted from oneend to the opposite end of the osteotomies (FIG. 21 c), consisting ofbirefringent, lamellar-fibered, compacta-type bone with a moderatenumber of vessels within the Haversian system (data not shown),characteristic of mature bone. In those instances in which there was buta partial bridging of the defect, there was often endochondralossification of cartilaginous islets. To exemplify, FIG. 22 aillustrates a typical endochondral cap at the medial aspect of theregeneration front: Hypertrophic chondrocytes were focally present inthe cartilaginous cap, which was enclosed by a thin, perichondrium-likefibrous tissue with parallel-oriented mature fibrocytes at its leadingedge (FIG. 22 b).

Hydrogel Degradadon and Osteogenesis—The extent as well as thedistribution of the osteoneogenesis depended on the erosion pattern ofthe hydrogel material and its relative placement in the defect siteafter the 5-week follow-up. Most notably, in treatment-2 rats, theregenerated bone extended well into the defect site, which includedresidual hydrogel surrounded by either a fibrous capsule or a foreignbody-type granulation tissue (FIGS. 21 a-c). Also, fibro-fatty tissueprimarily composed of fat cells with a minor fibrous component orfibrous tissue with a minor mononuclear celled inflammatory infiltrateoccurred in between the degraded hydrogel and the front of the newlyformed bone (FIG. 22 a). In contrast to the treatment-2 rats, thecharacteristic features of the fast-degrading hydrogels in the implantsite (treatment-1) resembled a typical nonunion with loosely organizedand edematous fibrous tissue with a minor non-specific chronicinflammatory infiltrate (FIGS. 5 a-c). The well vascularized connectivetissue within the gap was typical of that described in the literature asbeing consistent with nonunions, the distinctive picture which weencountered in all the control rats. Animals which did not show anysigns of new bone formation in the site of the defect exhibitedlongitudinally oriented myofibers which, so it seems, passively fellinto the defect, particularly at its center. There was subperiosteal newbone formation at the end of the long bone segment in some cases. Thedefect site was filled with the hydrogel surrounded by a fibrous capsuleand a subperiosteally regenerated bone at the osteotomy locales in thoserats treated with the slow degrading hydrogels (treatment-3).

Implant Histocompatibility: The PEG-fibrinogen hydrogels were alsoevaluated for histo(in)compatibility within the site of the tibialdefect. Microscopically, hydrogels which were partially degraded butremained in the defect after 5 weeks (treatment-2) revealed a serpentineforeign body type granulation tissue nearby residues of thePEG-fibrinogen hydrogels (FIG. 24 a). This granulation consisted oflymphocytes admixed with macrophages. At the tissue-implant interfacethe hydrogel was undergoing erosion by the finger-like projections ofthe granulation tissue that degraded the material and cleared it fromthe surface via the phagocytic pathway (FIG. 24 b). Encircling theserpentine projections of the above mentioned granulation tissue was achronic, non-specific inflammatory infiltrate consisting mainly oflymphocytes with a minor component of macrophages. The hydrogel itselfappeared diffusely homogeneous and was lightly colored in theH&E-stained sections; it was noteworthy that there was no cellularinfiltrate apart from that at the eroding front of the granulationtissue (FIG. 24 a, solid arrows). The response was restricted to just achronic, non-specific inflammatory reaction in certain regions of thetissue-material interface whereas elsewhere there had evolved apallisading foreign body type granulation tissue (FIG. 24 a, dashedarrows). The infiltration was focally admixed with a major infiltrate ofneutrophiles, while at the same time those implants which had not beendegraded after 5 weeks (treatment-3) were generally enclosed within afibrous capsule composed of several layer of elongated parallel orientedfibrocytes (FIG. 23 b). It thus stood to reason that the aforementionedgranulomatous response to the implanted hydrogel was indicative of thehistoincompatibility of the fibrin-based materials as described indetail [Boss J H. Biocompatibility: Review of the concept and itsrelevance to clinical practice. In: Wise D L, editor. Biomaterials andBioengineering Handbook. New York, Basel: Marcel Dekker, Inc.; 2000. p.1-94] in as much as there was a foreign body type response nearby theimplant.

Analysis and Discussion

This study shows that the resilience of the PEG-fibrinogen matrix can besynchronized with the optimal healing characteristics of a site-specificbone defect. Hydrogels having three distinctive compositions designedfor slow, intermediate, and fast degradation rates were fabricated andcharacterized (Table 4). The hydrogels were implanted into site-specificdefects of the tibiae of rats; the rationale being that the ingrowthmatrix would displace the post traumatic fibrin clot while sustaining asimilar healing effect in the site of the defect for a longer duration.There is undisputable evidence of extensive osteoneogenesis within thesite of the osseous defect based on recorded radiographic andhistological findings at the 5-weeks postoperative interval (FIGS. 21a-c). Moreover, the extent of newly formed bone within the gap is wellcorrelated to the degradation of the hydrogel matrix. Nevertheless, theprecise process by which osteoneogenesis occurs with an erodingPEG-fibrinogen hydrogel in the defect remains unclear.

In many histopathology sections of the treatment-2 rats residues of thehydrogel were detected apparently giving way to regenerating bonepenetrating the boney defect from both the proximal or distal aspects ofthe osteotomies. It may be concluded, therefore, that the hydrogelundergoing erosion clears the way for the regenerating bone (FIG. 22 a).The hydrogel in some cases has been completely degraded such that thewhole osseous defect has been filled with newly deposited bone from oneside of the osteotomy to the other side (FIG. 21 c). In contrast to thetreatment-2 rats, remnants of fast degrading hydrogels (treatment-1)have neither been observed within the site of the defect nor hasosteoneogenesis occurred within the site-specific defect of these rats(FIG. 23 a). It is, therefore, likely that these hydrogels underwentrapid proteolysis in the site of the osseous defects and were notpresent at the 5 week postoperative interval. Consequently, all thetreatment-1 animals exhibited extensive fibrotic scaring in the gap(FIG. 23 a). It may be speculated in this context that treatment-1 ratsbehave in a similar fashion to the non-treated control rats in that thefibrin-like clot disappears from the gap early on in the healing processand does not provide the necessary protection from the invadingfibrocytes as is characteristic in cases of nonunions. Conversely, theslow-degrading hydrogels (treatment-3) remain intact in the defect sitefollowing 5 weeks and exhibit only a local osteoneogenesis nearby theimplant (FIG. 23 b).

In the course of these experiments it has become apparent that only theintermediate degrading PEG-fibrinogen hydrogel treatment (treatment-2)causes extensive new bone formation at the site of the boney defect.Nevertheless, judging from just the histological findings it is unclearwhether the PEGylated fibrinogen material is endowed with osteoinductiveproperties or whether the so-called synchronized degradation and healingresponse are responsible for the widespread osteoneogenesis. There areat least a couple of plausible hypotheses that could explain theobserved outcomes in the above described experimental setup. Firstly, itis possible that the macrophages that erode the fibrin-basedbiosynthetic matrix slowly release osteoinductive fragments of thefibrinogen to act as an eroding front for osteoneogenesis to occur in asmuch as the gels slowly give way for the newly generating bone.Secondly, subperiosteal new bone is deposited by the osteoblastsconcurrently with the removal of the hydrogel by the foreign bodyinduced granulation tissue. In the case of the second option, thePEG-fibrinogen does not necessarily need to possess osteoinductiveproperties to facilitate the observed healing response within thesite-specific defect. Both explanations are consistent with theobservations that faster degrading hydrogels do not provide synchronizederosion with the natural healing rate in a rat site-specific bonedefect, which typically requires 4 to 5 weeks to heal completely.

The current inventors propose that the fragments released from thePEG-fibrinogen hydrogel probably facilitate a prolonged osteogenicresponse within the site-specific defect, thereby explaining the unusualextent of the oesteogenic response in the treatment-2 rats. There isample evidence that fibrinogen and fibrin degradation products arepotent agonist of wound healing [Thompson W D, et al., J Pathol1991;165(4):311-8; Rybarczyk B J, et al., Blood 2003;102(12):4035-43],especially as concerns endothelial cells [Lorenzet R, et al., ThrombHaemost 1992;68(3):357-63; Bootle-Wilbraham C A, et al. Angiogenesis2001;4(4):269-75] and fibroblasts [Gray A J, et al; Am J Respir Cell MolBiol 1995;12(6):684-90; Gray A J, et al., J Cell Sci 1993;104 (Pt2):409-13]. In fact, fibrin has been evidenced to induce an osteogenicresponse in bone defects filed with osteoconductive materials [AbiramanS. et al., Biomaterials 2002;23(14):3023-31; Kania R E, et al., J BiomedMater Res 1998;43(1):38-45; Gray A J, et al., J Cell Sci 1993;104 (Pt2):409-13]. Moreover, if fibrinogen fragments do not possessosteoinductive qualities, a similar outcome to that demonstrated byother researchers who utilize the inert biomimetic ingrowth matriceswithout added growth factors would be expected. Pratt et al. havereported that eroding fibrin-mimetic hydrogels are unable to support newbone formation when occupying a size-specific calvarial defect in theabsence of the osteoinductive BMP-2 [Pratt A B, Biotechnol Bioeng2004;86(1):27-36]. Likewise, Lutolf et al. have demonstrated similarresults employing a collagen-mimetic biosynthetic ingrowth matrixwithout osteoinductive BMP-2 [Lutolf M P, et al. Nat Biotechnol2003;21(5):513-8]. In comparison, the extent of osteoneogenesis usingPEG-fibrinogen hydrogels without added osteoinductive growth factors canonly be explained by an osteoinductive role of the fibrinogenconstituent in as much as it is unlikely that the PEG constituent hasosteoinductive qualities. In support of the proposal that protein-basedhydrogels illicit an osteoinductive response in the model of asite-specific bone defect, Ikada and co-workers have demonstratedosteoneogenesis in a calvarial defect filled with a gelatin ingrowthmatrix in the absence of osteoinductive growth factors, though only whenthe hydrogels are of a high water content [Hong L, et al. J Neurosurg2000;92(2):315-25; Yamamoto M, et al. J Control Release2000;64(1-3):133-42].

There is some discrepancy between the mostly mild osteogenic potentialof FSs as reported in the relevant literature and the extensiveosteoneogenesis discerned in the current study. The observedosteoneogenesis may be attributable to the released fragments ofPEGylated fibrinogen and not necessarily to the intact matrix. Asustained presentation of mildly osteogenic fibrinogen fragments couldaccount for the prolonged osteogenic response over the 5 weeks of thehealing period. It is noted in support of this explanation that most ofthe osteoneogenesis in the treatment-2 defects occurs at least severalhundred microns from the eroding surface of the hydrogels (FIG. 24 a),the latter being consistently surrounded by an inflammatory infiltrate.Even the slow degrading hydrogels (treatment-3) induce some mildosteogenic response around the implant, presumably because of thereleased fragments of PEGylated fibrinogen. There is no evidence ofosteoneogenesis in the fast degrading hydrogel-treated animals(treatment-1), suggesting that rapid dissolution of the fibrinogenfragments does not enable adequate new bone formation.

Others authors have also commented on the importance of sustainedrelease of osteoinductive factors in mediating osteogenesis in bonydefects. Most notably, Yamamoto et al. reported on the synchronizedrelease of immobilized basic fibroblast growth factor (bFGF) frombiodegradable gelatin hydrogels in a rabbit calvarial defect model[Yamamoto M, et al. J Control Release 2000;64(1-3):13342]. Hong et al.utilized a similar concept with transforming growth factorbeta-1(TGF-beta1) [Hong L, et al. J Neurosurg 2000;92(2):315-25]. Inboth these studies, the results indicated that the sustained release ofeither bFGF or TGF-beta1 from the hydrogel with synchronizeddegradability is necessary to fully capitalize on the osteoinductiveproperties of one of the other of these factors. The factor retention inthe site of the defect is too limited to induce osteoneogenesis whenhydrogels that degrade too rapidly are employed. On the other hand, thematrix serves as a physical barrier for factor-induced bone regenerationin the calvarial defect when hydrogels that degrade to slowly are used.

In conclusion, it may be submitted that the efficacy of the osteogenicresponse described in this study evolves in the wake of the combinationof the synchronized degradation of the hydrogel and the sustainedrelease of fragment of osteoinductive fibrinogen. The PEGylatedfibrinogen matrix may, therefore, constitute a highly efficacious toolfor orthopaedic surgeons who are faced with the problematic task ofpromoting the healing of site-specific defects in patients with bonefractures complicated by obstinate nonunions.

Example 6 Controlling Three-Dimensional Neurite Outgrowth UsingPEG-Fibrinogen Hydrogels

To test the potential of the PEG-fibrinogen scaffold material tofacilitate nerve regeneration, a chicken embryo dorsal root ganglion(DRG) outgrowth model was used. In the initial stage of nerveregeneration fibrin provides Schwann cells environmental cues forproliferation, thus ensuring that there are enough cells to associatewith the regenerating neurons. In the absence of fibrin, Schwann cellscan then differentiate and re-myelinate the newly formed axons.Accordingly, this model implies that the untimely persistence of fibrinin the injury site can interfere with the delicate timing of the nerveregeneration process and disrupt the construction of functional nervetissue. Therefore the ability to control the degradation and removal ofthe fibrin matrix is crucial for enabling successful nerve regeneration.

The synthetic PEG component provides the desired physical properties andcontrollable degradation characteristics. The natural fibrinogencomponent of the biosynthetic matrix supplies cues that regulate Schwanncell proliferation and migration and therefore will likely influencere-myelination of the regenerated axons. An additional advantage of thePEGylated fibrinogen approach is that it enables the control of therelative bioactivity of the fibrinogen degradation products based on therationale that covalenily bound PEG can decrease the accessibility toactive sites on both intact and degraded fibrinogen molecule. Hence, aPEGylation strategy offers control over fibrinogen degradation,bioactivity, and molecular architecture of the nerve guidance conduit(NGC) cell ingrowth matrix.

In addition, the present invenntors have shown that it is possible tocontrol the biodegradation of the fibrinogen matrix by changing relativeamounts of fibrinogen and PEG in such a system. To this end, Dikovsky etal. showed that increased PEG-DA concentrations in the PEGylatedfibrinogen hydrogel decreased proteolytic susceptibility of the proteinbackbone and thus delayed the PEGylated fibrinogen biodegradation[Dikovsky D. et al. Biomaterials 2006;27(8):1496-506]. The PEGylatedfibrinogen system also presents additional advantages for nerveregeneration in that therapeutic growth factors can easily beencapsulated and enmeshed in the dense polymeric network of the hydrogelduring the polymerization process. The encapsulation of factors fornerve regeneration could provide neuron-specific signals beyond theinherent biological and structural provisions of the PEGylatedfibrinogen hydrogel network. One of the most vital neurotrophins inneuronal development and regeneration is nerve growth factor (NGF).Schwann cells produce NGF, a 26 kDa dimmer, in their immature phaseduring early development and after post-injury dedifferentiation inmature nerves. Accordingly, it is important that NGF be an integral partof the nerve guidance implant material.

Materials and Experimental Methods

Dorsal Root Ganglia Experiments: DRGs were dissected from E9-E11 chickenembryos and collected in PBS with 1% penicillin-streptomycin (Biologicalindustries, Kibbutz Beit Haemek, Israel). Fibroblast contamination ofthe DRGs was minimized by pre-plating the DRGs for one hour in MEM withGlutamax I medium (Gibco, Grand Island, N.Y., USA) containing 1%penicillin-streptomycin and 10% fetal calf serum (FCS) (Biologicalindustries, Kibbutz Beit Haemek, Israel). The pre-plated DRGs were thenphysically removed from the culture dish and entrapped in hydrogelconstructs prepared from a precursor solution of PEGylated fibrinogen(prepared as described in Example 1) and photoinitiator. Briefly, theprecursor solution was mixed with 1% (v/v) photoinitiator stock solutionmade of 10% (w/v) Irgacure™2959 (Ciba Specialty Chemicals, Tarrytwon,N.Y.) in 70% ethanol and deionized water. The solution was thencentrifuged at 14,000 RPM for 1 minute before being used to entrap theisolated DRGs. The entrapment procedure involved gently placing theintact DRGs into a 48-well plate containing the precursor solution. The48-well plate was first pre-coated with 100 μl polymerized PEGylatedfibrinogen in order to prevent cell growth on the bottom of the well.Each DRG was placed into 200 μl PEGylated fibrinogen solution andpolymerized under a UV light (365 nm, 4-5 mW/cm²) for 5 minutes. Afterhydrogel polymerization, the entrapped DRGs were visually inspected toensure 3-D encapsulation in the biosynthetic matrix (FIGS. 25 a-c).Culture medium was immediately added to the polymerized hydrogels (500μl in each well) and changed every two days. The culture medium wascomprised of MEM with Glutamax I medium containing 1%penicillin-streptomycin and 10% FCS. Unless otherwise indicated, themedium was supplemented with 50 ng/ml 2.5S mouse nerve growth factor(mNGF) (Alomone labs LTD., Jerusalem, Israel).

Quantitative Outgrowth Measurements: Cellular outgrowth from the DRGinto the transparent PEGylated fibrinogen hydrogel was recorded duringthe four-day duration of the experiment. Each DRG construct wasdocumented with digital images taken daily using a Nikon TE2000 phasecontrast microscope with a 4× objective and a digital CCD camera(Jenoptik, Germany). Quantitative neurite outgrowth measurements wereobtained directly from the digital phase contrast micrographs usingImageJ software. Neurites, which can be identified by theircharacteristic sprouting morphology, were measured from the base (outermargin of the DRG) along their length and up to the tip. Up to a totalof 80 measurements were made for each DRG construct, according to theability to trace continuous neurites. The mean DRG neurite outgrowth wasthen calculated for each individual DRG construct by averaging the 80measurements of each construct (n=1). The average neurite outgrowth foreach treatment was calculated using the mean DRG neurite outgrowth data.

Histology and Immunofluorescence: Preparation of the DRG specimens forhistological and immunofluorescence evaluation involved fixation in 4%paraformaldehyde (Gadot, Haifa, Israel) for 20-30 min, PBS rinses, andovernight cryoprotection in a 30% sucrose solution (in PBS) at 4° C.Each fixed construct was then slow-frozen in Tissue-Tek® O.C.T Compound(Sakura Finetek, Torrance, Calif., USA) using liquid nitrogen cooledisopropanol (Gadot, Haifa, Israel). Frozen constructs were stored in adeep freezer (−80° C.) for up to three months. The specimens weresectioned orthogonally into 30-μm thick slices using a cryostat andmounted on Polysin™ slides (Menzel-Glaser, Braunschweig, Germany). Priorto staining, the slides were air dried at RT for 2 hours and stored at−20° C. Hematoxylin and Eosin (H&E) staining (Sigma, St. Louis, Mo.,USA) was performed according to standard manufacturer's protocols.

Immunofluoerscence labeling of the 30-μm thick specimens involvedtreatment with 0.3% Triton® X-100 (Bio Lab LTD., Jerusalem, Israel) for5 min at RT and incubation in blocking solution containing PBS and 1%glycine, 10% horse donor serum (HDS) (Biological industries, KibbutzBeit Haemek, Israel) and 0.1% Triton® X-100 for 30 min at RT. Thesections were double stained with primary antibodies againstβIII-tubulin, (G712A, Promega, Madison, Wis., USA) and s100 (S2644,Sigma, St. Louis, Mo., USA). The primary antibodies were diluted inblocking solution (1:1000 dilution for βIII-tubulin and 1:200 dilutionfor s100) and incubated overnight at 4° C. in a humidity chamber. Thesections were rinsed and incubated for 30 minutes at RT withfluorescently conjugated secondary antibodies, including 1:250 dilutedgoat anti-mouse Cy3 (Chemicon International, Temecula, Calif., USA) forβIII-tubulin and 1:300 diluted goat anti-rabbit FITC (JacksonImmunoresearch Laboratories INC., west Grove, Pa., USA) for s100. Anuclear counter-stain was incorporated directly into the secondaryantibody staining solution using a 1:500 diluted DAPI stock solution(Sigma Aldrich, St. Louis, Mo., USA). Following incubation, sectionswere rinsed with PBS and mounted with FluoromountG (SouthernBiotechnology Associates, INC., Birmingham, Ala., USA).

Statistical analysis: Statistical analysis was preformed on data setsfrom at least two independent experiments. Depending on the data set,treatments were compared by single-factor ANOVA, two-factor ANOVA, orpaired student t-test. Statistical significance was accepted for p<0.01.

Experimental Results

DRG Outgrowth: Tissue constructs were prepared by entrapping DRGs insidePEGylated fibrinogen hydrogels (FIGS. 25 a-c) and cultivating them forup to one month in a CO₂ incubator. Cellular outgrowth from the DRG wasvisible in phase contrast micrographs and histological H&E sections(FIGS. 26 a-d). Throughout the experiment, cells from the DRG invadedthe PEGylated fibrinogen hydrogel and eventually occupied the entire gel(not shown). Phase contrast micrographs show the distinct spatialorganization and orientation of the invading DRG cells into thePEGylated fibrinogen matrix after two days (FIG. 26 a). A highmagnification of this organization is shown in FIG. 26 b, where longthin processes (neurites) extending out of the DRG are accompanied bynon-neuronal cells (dark circular spots) that emerge from the DRG coreand align along the neurite extensions. The non-neuronal outgrowth fromDRGs (FIG. 26 b, arrowhead) was shown to lag after neurite extensions(FIG. 26 b, arrow). Histological cross-sections (30 μm) of the DRGconstructs following four days of culture stained with H&E showedsimilar cellular invasion characteristics (FIGS. 26 c-d).

The arrangement of non-neuronal (glial) cells invading from the DRG andaligning with the neurites resembled the in vivo spatial organization ofneurons and their associated Schwann cells. In order to identify thedifferent invading DRG cells in these experiments, neurites and Schwanncells were both labeled with neuronal and glial immunofluorescentmarkers. Immuno-detection in 30 μm thick cross-sections of the DRGconstructs was preformed with the neuronal marker βIII-tubulin antibodyand the Schwann cell marker s100 antibody. The labeling clearly showsextending neurites originating from the DRG into the matrix, andassociated Schwann cells in close proximity to the invading neuronalcells (FIGS. 27 a-c). Higher magnification images show the Schwann cellsclosely associated with the neurites to the extent that they align alongwith and adjacent to the βIII-tubulin positive extensions (FIGS. 27d-f). These results were well correlated to observations of the DRGcells inside the hydrogel as observed by phase contrast microscopy(FIGS. 26 a-b).

Nerve Growth Factor Treatments: Experiments to examine the influence ofNGF in the culture medium versus encapsulated in the hydrogel during itsformation were performed with DRG outgrowth constructs. Three treatmentconditions were compared: a treatment using no NGF (NO-NGF), a treatmentusing free-soluble NGF in the culture medium (FS-NGF), and a treatmentwith enmeshed NGF in the hydrogel network (EN-NGF). Two independentexperiments in each treatment condition were preformed for a total ofsix repeats using two different batches of PEGylated fibrinogenprecursors. The constructs were cultured for four days and imaged dailyto measure the progress of 3-D cell outgrowth from the DRGs into thehydrogel network. Based on results from the phase contrast micrographs(data not shown), the free-soluble and enmeshed NGF (FS-NGF and EN-NGF)facilitated outgrowth of both non-neuronal cells and neurites into thehydrogel as compared to NGF-deprived constructs (NO-NGF). In the absenceof NGF, there was no observable outgrowth of neurites and only partialoutgrowth of non-neuronal cells, which were most likely Schwann cells orfibroblasts. Immunohistochemistry confirms the observations of phasecontrast microscopy in that βIII-tubulin and s100 positive cells werepresent in NGF treatments (FS-NGF and EN-NGF) but only s100 positivecells were seen in the NGF-deprived treatment (NO-NGF) (FIGS. 28 a-c).Based on these qualitative data, it is difficult to conclude if thereare significant differences in 3-D DRG outgrowth between the freesoluble and enmeshed NGF; both free soluble NGF (FS-NGF) and enmeshedNGF (EN-NGF) treatments showed a similar labeling pattern.

In order to further differentiate between the free soluble and enmeshedNGF treatments, quantitative outgrowth experiments were performed. Usingdigital image processing, the distance of neurite outgrowth was measuredin DRG constructs that were cultured with free soluble NGF (FS-NGF) orenmeshed NGF (EN-NGF). FIG. 28 d shows that there was little differencebetween the two treatment conditions at any time during the cultureperiod (p>0.35, n=6). In both free soluble and enmeshed NGF, there was arapid increase in neurite outgrowth over the course of the four-dayexperiment (P<0.01, n=6), with a mean neurite length reaching 719.9 μmand 701.2 μm for FS-NGF and EN-NGF treatments, respectively after fourdays.

Cellular Outgrowth and Hydrogel Biodegradation: Alterations to thebiodegradation properties of the fibrinogen backbone of the hydrogelmatrix can also influence the DRG cellular outgrowth characteristics,particularly as related to the relative invasion of Schwann cells andneurites. Experiments were performed to assess the ability to regulatethe outgrowth kinetics using different compositions of the matrix(relative amount of PEG and fibrinogen) based on the rationale that theproteolytic resistance of the fibrinogen matrix will increase withincreasing concentrations of PEG. Consequently, the PEGylated fibrinogenhydrogels also become more cross-linked with additional PEG, therebychanging the mesh size, hydration and mechanical properties of thematrix. Four different compositions of PEG to fibrinogen were tested,including: 30:1, 60:1, 120:1, and 180:1 (PEG:fibrinogen). It isimportant to note that the composition of the constructs in eachtreatment level was such that the pure PEGylated fibrinogen solution(30:1 treatment) was modified with additional unreacted PEG-DA beforethe UV polymerization step. Two independent experiments in eachtreatment level were preformed for a total of nine repeats using twodifferent batches of pure PEGylated fibrinogen precursors.

Overall, the extent of cellular outgrowth from the DRG into the matrixwas decreased with addition of higher concentrations of PEG-DA in thehydrogel matrix (FIGS. 29 a-p). The lag between neurites and glial cellswas visibly reduced with the addition of higher concentrations of PEG-DA(FIGS. 29 q-t). A summary of the neurite outgrowth kinetics data withthe different concentrations of PEG is summarized in FIG. 29 u.Statistical analysis of the kinetics data (2-factor ANOVA) revealed thatoutgrowth steadily increased with culture time (p<0.01, n=9) and thathigher concentrations of PEG slowed down the cellular invasion (p<0.01,n=9). In particular, constructs made with high concentrations of PEG-DA(120:1 and 180:1) delayed neurite outgrowth significantly from day 2 ofculture when compared to constructs made with lower concentrations ofPEG (30:1 and 60:1). Neurite outgrowth in the 180:1 hydrogels exhibitedslowest outgrowth kinetics of all treatment levels, and reached a meanneurite length of 201.6 μm following four days of culture. Constructsmade with 120:1 exhibited moderate neurite outgrowth rate and reached amean neurite length of 490.8 μm following four days. There was nosignificant difference in neurite outgrowth between the 30:1 and 60:1treatment levels (p>0.50, n=9); in both cases, outgrowth progressed mostrapidly and reached a mean neurite length of 807.8 μm and 850.6 μm,respectively following four days of culture.

Fibrinogen and DRG Cellular Outgrowth: The importance of the fibrinogenbackbone in enabling cellular outgrowth from the DRG into the PEGylatedfibrinogen matrix was investigated using PEG-only hydrogels as controls.DRG constructs were made of 10% PEG-DA without fibrinogen and comparedto constructs made with PEGylated fibrinogen. The constructs werecultured for three days and cellular outgrowth was documented on thethird day of culture. FIG. 30 a shows that without fibrinogen, very fewneurites extend out of the DRG and outgrowth of non-neuronal cells,including Schwann cells, was not observed. In contrast, fibrinogencontaining hydrogels exhibit massive DRG outgrowth, including neuriteand Schwann cell invasion, following three days of culture (FIG. 30 b).These results demonstrate fibrinogen's role in permitting DRG outgrowththat includes proteolytic susceptibility, inductive and conductiveenvironmental cues which may be crucial for functional peripheral nerveregeneration. Consequently, neuronal outgrowth was practicallyeliminated even in the PEGylated fibrinogen hydrogels when DRG cultureswere deprived of NGF (NO-NGF), whereas other cell types (includingSchwann cells) are observed invading the hydrogel (FIG. 30 c).

Analysis and Discussion

Peripheral nerve regeneration is a complex, highly regulated processwhich requires specific environmental cues that are provided by theextracellular matrix (ECM) and the tight bi-directional communicationbetween regenerating axons and their associated Schwann cells. Manyperipheral nerve regeneration strategies using NGCs have been designedto provide the optimal milieu for PNS regeneration, using natural orsynthetic materials and different growth factor delivery strategies.Because NGCs have yet to achieve the efficacy of the nerve autografts,alternative approaches are sought that can leverage the natural healingmechanisms of peripheral nerve repair following moderate injury. To thisend, the PEGylated protein hydrogels of the present invention can serveas a template for nerve regeneration which combines the paracrineeffects of fibrin(ogen) and the control over biodegradation andbioactivity afforded by the PEGylation paradigm.

The experiments descibed hereinabove support a potential NGC biomaterialsystem based on PEGylated fibrinogen hydrogels that maintain outgrowthof DRG cells. A 3-D hydrogel matrix composed of PEG and fibrinogen wasused to encapsulate chicken embryo DRGs to form transparent constructsthat enable straightforward monitoring of the DRG outgrowth (FIGS. 25a-c). Outgrowth of neurites and non-neuronal (glial) cells was observedfrom the DRG into the hydrogel (FIGS. 26 a-d). Furthermore, in vivo likespatial organization of these cells was observed. Specifically, the longneurites were observed in close proximity to their associated glialcells. Using antibodies specific for βIII-tubulin and s100, it was shownthat the s100-positive Schwann cells are highly associated with theradially extending βIII-tubulin positive neurites (FIG. 27 a-f). Theseneuron-Schwann cell complexes enable the production and organization ofmyelin along the length of extending axons in order to provide rapid andefficient propagation of action potentials along axons. Consequently,this distinct spatial organization is a prerequisite for axonalmyelination during the later phase of neuronal regeneration.

It is likely that DRG neurites and glial cells employ a proteolyticmechanism to invade the PEGylated fibrinogen hydrogel matrix in as muchas the hydrogel is highly susceptible to proteases [Almany L, et al.Biomaterials 2005;26(15):2467-77] and is otherwise too dense to permitcellular invasion in the absence of proteolysis. Because the fibrinogenbackbone affords the biosynthetic hydrogel its biodegradability, it alsoprovides a means of releasing cleaved fragments of fibrinogen from thematrix upon degradation. In this manner the kinetics of neurite andglial cell invasion as well as the bioactivity of the releasedfibrinogen fragments can be controlled by changing the relative amountof PEG and fibrinogen. Higher amounts of PEG reduce the susceptibilityto proteolytic degradation of the fibrinogen backbone [Dikovsky D, etal, Biomaterials 2006;27(8):1496-506] and presumably reduces the overallbioactivity of the degraded fibrinogen fragments that are released tothe surrounding tissue [Hooftman G, et al. J Bioact Compat Polym1996;11:135-159]. Indeed, the addition of PEG to the biosynthetichydrogel slows down the invasion of both Schwann cells and neurites fromthe DRG (FIGS. 29 a-p). Furthermore, it appeared that in lowerconcentrations of PEG, the non-neuronal outgrowth from DRGs laggedbehind the neurite extensions, whereas the higher concentrations of PEGminimized this lag (FIGS. 29 q-t).

It is appreciated that certain features of the invention, which are, forclarity, described in the context of separate embodiments, may also beprovided in combination in a single embodiment. Conversely, variousfeatures of the invention, which are, for brevity, described in thecontext of a single embodiment, may also be provided separately or inany suitable subcombination.

Although the invention has been described in conjunction with specificembodiments thereof, it is evident that many alternatives, modificationsand variations will be apparent to those skilled in the art.Accordingly, it is intended to embrace all such alternatives,modifications and variations that fall within the spirit and broad scopeof the appended claims. All publications, patents and patentapplications mentioned in this specification are herein incorporated intheir entirety by reference into the specification, to the same extentas if each individual publication, patent or patent application wasspecifically and individually indicated to be incorporated herein byreference. In addition, citation or identification of any reference inthis application shall not be construed as an admission that suchreference is available as prior art to the present invention.

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1. A precursor molecule comprising a fibrinogen protein which isdenatured and retains an activity of forming a scaffold or a collagenprotein which is denatured and retains an activity of forming a scaffoldand at least two PEG molecules covalently connected to free thiol groupsof said denatured fibrinogen protein or said collagen protein, each ofsaid at least two PEG molecules comprising a functional group forcross-linking.
 2. The precursor molecule of claim 1, wherein said PEG isselected from the group consisting of PEG-acrylate (PEG-Ac) andPEG-vinylsulfone (PEG-VS).
 3. The precursor molecule of claim 2, whereinsaid PEG-Ac is selected from the group consisting of PEG-DA, 4-arm starPEG multi-Acrylate and 8-arm star PEG multi-Acrylate.
 4. The precursormolecule of claim 3, wherein said PEG-DA is a 4-kDa PEG-DA, 6-kDaPEG-DA, 10-kDa PEG-DA, 14-kDa PEG-DA and/or 20-kDa PEG-DA.
 5. Theprecursor molecule of claim 3, wherein a molar ratio ofPEG-DA:fibrinogen is 2-400:
 1. 6. A method of generating a scaffoldcomprising: (a) generating a plurality of precursor molecules, whereineach of said precursor molecules comprise a fibrinogen protein which isdenatured and retains an activity of forming a scaffold or a collagenprotein which is denatured and retains an activity of forming a scaffoldand at least two PEG molecules covalently connected to free thiol groupsof said denatured fibrinogen protein or said collagen protein, each ofsaid at least two PEG molecules comprising a functional group forcross-linking; and subsequently (b) cross-linking said plurality ofprecursor molecules to thereby generate the scaffold.
 7. The precursormolecule of claim 1, wherein said fibrinogen or collagen protein retainsan activity of mediating tissue regeneration following in vivoadministration.
 8. A precursor molecule comprising a fibrinogen proteinwhich is denatured and retains an activity of mediating tissueregeneration following in vivo administration, or a collagen proteinwhich is denatured and retains an activity of mediating tissueregeneration following in vivo administration, and at least two PEGmolecules covalently connected to free thiol groups of said fibrinogenprotein or said collagen protein, each of said at least two PEGmolecules comprising a functional group for cross-linking.